Foam prosthesis for spinal disc

ABSTRACT

Disclosed herein are spinal disc implants comprising a foam adapted to completely or partially replace a nucleus pulposus within a spinal disc cavity, the foam being a nonabsorbable, closed cell and having a Poisson ratio of less than 0.5. Also disclosed are methods of implanting a foam, either as an in-situ curable material or as a preformed foam.

RELATED APPLICATIONS

This application claims the benefit of priority under 35 U.S.C. §119(e)of U.S. Provisional Application No. 60/931,407, filed May 22, 2007, U.S.Provisional Application No. 60/930,064, filed May 14, 2007, and U.S.Provisional Application No. 60/930,104, filed May 14, 2007, thedisclosures of which are incorporated herein by reference.

FIELD OF THE INVENTION

Disclosed herein are prostheses for replacing all or part of a nucleuspulposus in a spinal disc area.

BACKGROUND OF THE INVENTION

This disclosure relates to treatments and implant materials for treatingthe medical condition called degenerative disc disease (DDD) byaugmenting the intervertebral disc space, with or without the removal ofintradiscal tissue, the nucleus pulposus (NP), and localizing,supplementing, or replacing native NP with an in situ curing polymer(Nucleus Replacement Prosthetic or NRP) while the outer intervertebralmembrane called the annulus fibrosus (AF) heals after rupture andsurgical intervention commonly referred to as a discectomy procedure.The NRP is designed to treat back pain and be both physiologically andmechanically supportive of the pathological intervertebral disc (IVD) toultimately preserve the natural motion in the vertebral segment.

Treating DDD by performing a discectomy can include surgical accessthrough the AF and the removal of all or some the NP through an existingrupture in the AF or a surgical incision through the AF called anannulotomy. The annulotomy or disc rupture is a tunnel throughindividual fibro-cartilage rings, or laminae, that make up the AF.Typically the annulotomy/rupture is not closed after removal of the NPand other fibrous debris. The open annulotomy will scar over andeventually heal within weeks of the surgery. The medical literaturecites a 5% to 11% risk of re-operation (depending on the amount removalof the NP) to remove additional intradiscal debris that has migratedfrom the intradiscal space through the annulotomy and has come intocontact with sensitive spinal nerves around the surgical site before theannulotomy has scarred shut. It is often a clinical objective of thenucleus replacement prosthetic to localize any intradiscal debris afterpartial nuclectomy while the annulotomy scars shut, sealing theintradiscal space naturally.

Another complication of the discectomy procedure is the loss of discheight (space between two vertebral bodies at any given spine level)after surgery. Disc height loss is directly related to radiculopathy orleg pain due to nerve compression. The medical literature cites that 20%to 30% disc height loss is anticipated in the first 12 months after thediscectomy procedure resulting in 5% to 30% re-operation risk withinfive or six years after surgery. A second but much longer clinicalobjective is to maintain disc height over time with the replacement ofNP with NRP vs. not replacing with NRP. The ability to replenishexisting implanted NRP with new NRP using non-surgical techniques isdesired, thus allowing for revision of existing NRP and indefinitelymaintain the functionality of the IVD for the life of the patient.

The concept of treating an intervertebral disc abnormality by placing anin vivo forming implant in the intradiscal space normally occupied bydisc nucleus or has its origin at least as early as the 1962. An in vivoforming implant is any substance placed medically in the body that isintended to persist for a therapeutic interval, sometimes as long as theremaining life of a patient, which undergoes a change in its mechanicalcharacteristics, generally a phase transition. The phase transition canbe due to a condition within the body that is different from the storagecondition of the implant, for example a temperature or pH change. Thephase transition can be induced by combining ingredients prior toimplantation resulting in polymerization, precipitation, or otherwisehardening of the implant within the body. The phase transition can beinduced by an outside stimulus, such as photo-initiating light or otherextra-corporeal sources of energy. In most cases the phase transition isfrom a liquid to a solid, or from a low modulus solid to a highermodulus solid involving a structural change in the implant.

The concept of achieving an in vivo forming transition in an implant forthe purpose of treating an intervertebral disc abnormality by means ofin situ curing also has its origins as early as the 1980's. The in situcuring pathway frequently chosen is chemical polymerization of simplefunctional groups that result in the formation of extended polymernetworks within living tissue. The functional groups most frequentlycited can be arranged in three broad categories: 1) functionalizedpolyols, 2) acrylates, and 3) functionalized or crosslinked proteins.

The category of functionalized polyols includes reactions that result inthe formation polyurethanes, polyurea urethanes, and other chemistriesinvolving active NCO end groups. For example, Garcia French patent2,639,823 discloses as early as 1988 the use of an in situ polymerizingpolyurethane mixture delivered into a enclosing device into the nuclearspace of a spinal disc. Felt U.S. Pat. Nos. 5,888,220; 6,140,452;6,248,131; 6,306,177; 6,443,988; 6,652,587 and 7,001,431 (Bao et al)disclose a two-part in situ polymerizing polyurethane composition thatis injected into a molding device placed in the nuclear space of aspinal disc. Later, Bao and Felt U.S. Pat. No. 7,077,865 and U.S.published applications 2007/0038300 and 2006/0253200 disclose the use ofin situ curable compositions without the use of a molding device such asa balloon. Milbocker et al, U.S. published application 2005/0070913discloses a one-part polymerizing polyurethane composition that bonds tothe tissue of the nuclear disk space as it polymerizes in situ to fill adisc nucleus, repair a disc Annulus, or localize a nuclear prosthetic.

There have since been a number of disclosures directed to filling thenuclear space of a spinal disc with a variety of in situ curingmaterials. Haldimann U.S. Pat. No. 6,428,576 discloses the use of insitu curing materials without specification of the implant composition.Higham U.S. published application 2006/0255503 discloses an in situcuring disc nucleus implant containing a radio-opaque agent. Kim U.S.published application 2007/0005140 discloses delivery techniquesemploying multiple injection of in situ curing implant. Trieu U.S.published applications 20006/0200245; 2006/0089719; 2006/0064172;2006/0064171; 2005/0203206 disclose various combinations of in situpolymerizing implants in conjunction with nuclear prosthetics tolocalize a prosthetic in a spinal disc.

With respect to particular compositions, Collins U.S. publishedapplication 2006/0009851 and 2006/0009778 disclose in situ polymerizingprotein compositions for treatment of a spinal disc nucleus. Umit U.S.published application 2007/0093902 discloses other in situ polymerizingprotein compositions for treatment of a spinal disc nucleus. In thecategory of acrylate compositions of in situ curing nuclear implants,Mallupragada U.S. Pat. No. 7,183,369 and Milner U.S. Pat. No. 6,187,048are examples.

Without specificity to spinal applications, there are a variety ofinjectable biomaterials disclosed in issued patents including:cross-linkable silk elastin copolymer disclosed in Stedronsky U.S. Pat.No. 6,423,333, Capello U.S. Pat. No. 6,380,154, Ferrari U.S. Pat. No.6,355,776, Stedronsky U.S. Pat. No. 6,258,872, Ferrari U.S. Pat. No.6,184,348, Ferrari U.S. Pat. No. 6,140,072; Stedronsky U.S. Pat. No.6,033,654; Ferrari U.S. Pat. No. 6,018,030; Stedronsky U.S. Pat. No.6,015,474; Ferrari U.S. Pat. No. 5,830,713; Stedronsky U.S. Pat. No.5,817,303; Donofrio U.S. Pat. No. 5,808,012; Capello U.S. Pat. No.5,773,577; Capello U.S. Pat. No. 5,773,249; Ferrari U.S. Pat. No.5,770,697; Stedronsky U.S. Pat. No. 5,760,004; Donofrio U.S. Pat. No.5,723,588; Ferrari U.S. Pat. No. 5,641,648; Capello U.S. Pat. No.5,235,041; protein hydrogel described in Morse U.S. Pat. No. 5,318,524;Morse U.S. Pat. No. 5,259,971; Morse U.S. Pat. No. 5,219,328;polyurethane-filled balloons disclosed in Bao U.S. Pat. No. 7,077,865;Bao U.S. Pat. No. 7,001,431; Felt U.S. Pat. No. 6,306,177; Felt U.S.Pat. No. 6,248,131; Bao U.S. Pat. No. 6,224,630; collagen-PEG disclosedin Olsen U.S. Pat. No. 6,428,978; Olsen U.S. Pat. No. 6,413,742; RheeU.S. Pat. No. 6,323,278; Wallace U.S. Pat. No. 6,312,725; Sierra U.S.Pat. No. 6,277,394; Rhee U.S. Pat. No. 6,166,130; Berg U.S. Pat. No.6,165,489; Simonyi U.S. Pat. No. 6,123,687; Berg U.S. Pat. No.6,111,165; Sierra U.S. Pat. No. 6,110,484; Prior U.S. Pat. No.6,096,309; Rhee U.S. Pat. No. 6,051,648; Esposito U.S. Pat. No.5,997,811; Berg U.S. Pat. No. 5,962,648; Rhee U.S. Pat. No. 5,936,035;Rhee U.S. Pat. No. 5,874,500; chitosan disclosed in Chemte U.S. Pat. No.6,344,488; other polymers discussed in Boyd U.S. Pat. No. 7,004,945;Collins U.S. publication 2006/0004326; Collins U.S. publication2006/0009851; Milner U.S. Pat. No. 6,187,048; Daniell U.S. Pat. No.6,004,782; Urry U.S. Pat. No. 5,064,430; Urry U.S. Pat. No. 4,898,962;Urry U.S. Pat. No. 4,870,055; Urry U.S. Pat. No. 4,783,523; Urry U.S.Pat. No. 4,589,882; Urry U.S. Pat. No. 4,500,700; Urry U.S. Pat. No.4,474,851; Urry U.S. Pat. No. 4,187,852; Urry U.S. Pat. No. 4,132,746.

None of the treatments or compositions and associated surgical methodsof their use described above are entirely satisfactory from either abiocompatibility or efficacy perspective for formation of a nucleusimplant in situ. Accordingly, there remains a need for continueddevelopment of implant materials and treatment methods.

SUMMARY OF THE INVENTION

One embodiment provides a spinal disc implant comprising a foam adaptedto completely or partially replace a nucleus pulposus within a spinaldisc cavity, the foam being a nonabsorbable, closed cell and having aPoisson ratio of less than 0.5.

Another embodiment provides a method of repairing a defect in a spinaldisc space, comprising:

inserting a nonabsorbable, closed cell foam having a Poisson ratio ofless than 0.5 into the defect.

A method of repairing a defect in a spinal disc space, comprising:

inserting a composition in the area of the defect, the compositioncomprising:

-   -   (a) a prepolymer, and    -   (b) a foaming component; and

curing the composition to form a nonabsorbable, closed cell foam havinga Poisson ratio of less than 0.5.

BRIEF DESCRIPTION OF THE DRAWINGS

Various embodiments will be understood from the following description,the appended claims and the accompanying drawings, in which:

FIG. 1 is a schematic view of an incompressible substance subjected to aloaded along its z-axis;

FIG. 2A is a schematic view of a spinal disk disc in transverse crosssection;

FIG. 2B is a schematic view of a disc in lateral cross section;

FIG. 3 is a schematic view of a disc in transverse cross section;

FIG. 4 is a plot of disc deflection under load as a function of hoopstress ratios;

FIG. 5 is a plot of disc deflection as a function of a spring constants;

FIG. 6 is a schematic view of in lateral cross-section a spinal disc iscomprising bony endplates, annulus fibrosus, and nucleus pulposus;

FIG. 7A is a schematic view of a disc in transverse cross-section;

FIG. 7B is a schematic view of a disc in lateral cross-section;

FIG. 8 is a plot of disc height as a function of increasing nucleusradius;

FIG. 9 is a plot of disc height lost as a function of nucleus to discradius;

FIG. 10 is a plot of load as a function of nucleus to disc radius;

FIG. 11 is a plot of disc height lost as a function of nucleus to discradius;

FIG. 12 is a schematic view of a disc in lateral cross-section withinset 253 showing the layered structure of endplate 107 and inset 252showing the layered structure of annulus 108;

FIG. 13 is a plot of the ratio of nucleus load/annulus load as afunction of disc deflection;

FIG. 14 is a plot of load as a function of displacement;

FIGS. 15A and 15B are plots of implant modulus as a function of nucleusradius for a fixed r_(n)/r_(d)=0.7 (FIG. 15A), and unfixed r_(n)/r_(d)(FIG. 15B);

FIG. 16 is a plot of implant modulus as a function of load;

FIG. 17 is a plot of implant modulus as a function of deflection;

FIG. 18 is a plot of implant modulus as a function of load;

FIG. 19 is a plot of annulus pressure as a function of prosthetic size;

FIG. 20 is a schematic view of forces involved in prosthetic extrusionthrough the annulus;

FIGS. 21A-C schematically depict stresses acting on an elemental sliceof the prosthetic;

FIG. 22 is a plot of failure pressure as a function of modulus;

FIG. 23 is a plot of nuclear pressure as a function of modulus;

FIG. 24 is a plot of load failure as a function of modulus;

FIG. 25 is a plot of load failure as a function of impact velocity;

FIG. 26 is a plot for the conversion of modulus (y-axis) to durometer(x-axis) for polyurethanes;

FIG. 27 is a plot of load failure as a function of modulus;

FIG. 28 is a plot of load failure as a function of modulus;

FIG. 29A is a plot of nucleus pressure as a function of defect diameter;

FIG. 29B is a plot of load as a function of defect diameter;

FIG. 30 is a plot of nucleus pressure as a function of defect diameter;

FIG. 31 is a schematic view of flexion-extension properties of a discunder load;

FIG. 32 is a schematic view of the annulus and nucleus applying forcesto the endplate;

FIG. 33 is a schematic view of the tilt component and effect on the discheight;

FIG. 34 is a plot of failure pressure/modulus ratio as a function ofdefect diameter;

FIG. 35 is a plot of restorative force as a function of displacement;

FIG. 36 is a plot of total load as a function of displacement; and

FIG. 37 is an illustration of common pathologies of spinal discs in thelumbar region.

DETAILED DESCRIPTION

Disclosed herein are treatments and implant materials disclosed relatingto optimizing the function of replacement nucleus pulposus prostheticsplaced. The characteristics of a nucleus replacement prosthetic can bederived from physical properties defined by the biomechanical andbiological requirements for localizing and maintaining the prosthetic inthe disc space.

One embodiment provides a spinal disc implant comprising a foam adaptedto completely or partially replace a nucleus pulposus within a discnucleus space, the foam being a nonabsorbable, closed cell and having aPoisson ratio of less than 0.5.

In one embodiment, the implant that partially or wholly replaces thenucleus pulposus is delivered to and localized in a disc nucleus space.The disc nucleus space includes the nucleus pulposus and optionally theadjacent tissues, including the vertebral endplates and inner layers ofthe annulus fibrosus (e.g., the layers contacting or immediatelysurrounding the nucleus pulposus).

Implants with a Poisson ratio of 0.5 are practically incompressibleunder normal physiologic loads. With such implants, a volumetricdecrease in implant height results in an equal volumetric expansion ofimplant radius when approximating the implant as a cylinder. For animplant of a given modulus, there is a fixed relationship between theaxial forces applied to the implant, the radial forced applied to theannulus and the changes in implant height and radius. For example,implants with Poisson ratio 0.5 cannot change height without alsochanging radius, and this incompressibility of the implant directlycouples the annulus to the endplates for all load frequencies. Previousimplant materials include gels and liquids that have a Poisson ratio of0.5 and have the disadvantage of incompressibility to loads.

In contrast, implants with a Poisson ratio less than 0.5 has somefraction of its total volume in the compressible form of a gas, e.g.,bubbles or closed cells in a foam implant. The greater the fraction ofbubbles in the implant the lower the Poisson ratio. Accordingly, in oneembodiment, the implant comprises a foam, such as a closed-cell foam.The compressibility affords a looser coupling of the height and radialdimensions at high load frequencies, as opposed to the decrease inheight an equal volumetric radial expansion of incompressible implants.For example, for a compressible foam that experiences a sudden impactdoes not subject the surrounding tissue to an immediate expansion. Thiscompressibility feature may help preserve the annulus and/or endplatesby storing transitionally some of the impact energy of the load aspotential energy in the implant.

The foam can be a preformed foam having the recited properties, or acurable foam implanted as a liquid implant material and cured while inthe disc space. Accordingly, disclosed herein are preformed foams and insitu curing nucleus implant foams intended to therapeutically replace oraugment the natural disc nucleus space. One embodiment provides acurable implant composition, comprising:

-   -   (a) a prepolymer, and    -   (b) a foaming component.

The prepolymer and foaming component is described in greater detailbelow. When combined, the composition can be cured to form anonabsorbable, closed cell foam having a Poisson ratio of less than 0.5.

Curing with respect to a liquid or deformable nucleus implant implies aphase change of the implant from liquid to solid or deformable solid toless deformable solid. One embodiment provides a liquid prepolymer mixedwith water, saline or a therapeutic aqueous solution prepared outsidethe body and injected into a vertebral disc. The implantation can befacilitated by the mixing of liquid prepolymer and aqueous solution toprovide for a low viscosity liquid implant that can be injected througha needle. In procedures directed to correcting a diseased or bulgingvertebral disc, often a hole is made in the annulus in order to removesome or the entire disc nucleus. In one embodiment, the removed nucleus,whether completely or partially removed, is replaced with an implant. Insuch cases a liquid implant would flow out of the implantation site, anda curing step localizes the implant in the disc.

The natural nucleus is largely water, and diseased nucleus in some casesis characterized as being dehydrated. In one embodiment, a nucleusimplant contains a large fraction of water; yet possess the shaperetentive features of a solid. The amount of water in the cured implantdetermines its freestanding Young's modulus. The Young's modulus canrange from as little as 0.5 MPa to as great as 10 MPa. Generally, thelower the modulus the greater the fraction of axial forces translatedradially to the annulus. The minimum modulus of the cured implant can bedetermined by the size of a defect in the annulus. In the case where theimplant is injected through a small gauge needle, 20 G or more, and noother defect exists or is created in the annulus, the modulus can rangefrom 0.5 MPa to 1.0 MPa. The relationship between defect area andminimum implant modulus can be calculated as described in the Appendixbelow. Applying assumptions about the disease state of the disc and thelikely surgical intervention the following table summarizes the findingsof minimum implant modulus for different surgical interventions.

TABLE 1 Treatment Paradigms for Nucleus Replacement Disc Height IncreaseIndication Bonded Restored Nucleus Modulus Procedure Annulotomy ThinDisc Yes Yes No 0.5 MPa   Trans-axial No Lumbar Thin Disc Yes No No 1MPa Trans-axial No Lumbar Thin Disc No Yes No 3 MPa Trans- No Lumbarannulus Thin Disc No No No 6 MPa Trans- No Lumbar annulus Black Disc YesEither Yes 1 MPa annulotomy <2.5 mm Bulging Disc Either No No 2 MPaAnnulotomy <4.9 mm Bulging Disc Yes No No 1 MPa Annulotomy <4.9 mmBulging Disc Yes Yes No 0.25 MPa   Annulotomy <4.9 mm Permanently No NoNo 8 MPa Annulotomy Any size compressed Permanently No No Yes 6 MPaAnnulotomy <3.5 mm compressed Permanently Yes No No 1 MPa Annulotomy<3.5 mm compressed

The minimum implant moduli associated with the various surgicalinterventions of Table 1 are general guidelines for selecting an implantmodulus, and can be further refined by reference back to the details ofthe calculations in the Appendix. The cured modulus can be varied byvarying the ratio of prepolymer to aqueous solution at the time ofpreparation outside the body.

Another embodiment provides a cured implant comprised of solid, liquidand gaseous parts. This can be achieved with a solid polymer matrixcontaining both liquid and gaseous fractions. One example is a hydrogel,as described in U.S. application Ser. No. 10/020,331, published U.S.Pub. No. 2003/0135238, the disclosure of which is incorporated herein byreference. The gaseous and liquid components of the hydrogel foamimplant are in equilibrium, with the volume of the implant varying withthe applied load, wherein the ratio of the volumes of the liquid andgaseous components can vary according to the magnitude of the load. Thisfeature has a direct analogue in the natural disc, where the disc heightis typically less during periods of activity (high load) and expandsduring periods of sleep (low load). Gas in the implant can be forcedinto solution during protracted periods of load, reducing disc heightand nuclear volume. During periods of rest, the dissolved gas comes outof solution and can fill the bubbles in the hydrogel foam. The timescaleon which gas is forced into solution and then subsequently released backinto the implant as a gas can be similar to the timescale for theincrease in normal disc height during periods of rest. Thus, thisaccommodative effect is not responsive to high frequency changes inload, such as experienced while running.

In one embodiment, the water fraction of the implant is bound in theimplant such that pressures of the magnitude commonly encountered in thedisc do not result in water being driven out of the implant by pressure.In one embodiment, the implant contains loosely bound water. In thisembodiment, the water contained in the implant can exchange with waterin surrounding tissue through a statistical process, which is largelyunaffected by pressure since the pressure in the implant and thepressure in the tissue are approximately equal. Even in the case wherethe pressure in the implant exceeds that of the surrounding tissue,water will not be driven out of the implant if the water bound in thehydrogel is localized by hydrogen bonding. This kind of localization ofwater within a polymer is often associated with hydrogels, which possessan affinity for water. This affinity for water is also a characteristicof normal nucleus pulposus. In one embodiment, the cured implant retainsits water fraction through hydrogen bonding

One embodiment provides a minimally invasive means of delivery through asmall access hole or needle injection and a second localization featurethat prevents the implant from migrating through the access hole andaway from the intended implantation site. The localization feature isthe result of the cured implant being larger than the access hole. Acurable nucleus implant can undergo two phase-transitions and alsoretain its original liquid phase forming a foamed hydrogel. Acompressible preformed foam implant can retain its original shape oncedelivered through the small access hole and inserted in the disc nucleusspace.

In another embodiment, the treatment involves delivery of an in situpolymerizing tissue adhesive into the treatment area of the nucleus ornuclear space. The implant, if an in-situ curing implant is delivered influid form typically down an access port leading to a space in the discwhere formerly nucleus material resided and now air resides. For apreformed foam, the implant is delivered in compressed form and allowedto expand once inside the nuclear disc space. These procedures aim tofill a void created surgically in the disc. To fill a surgically createdvoid, the implant can either displace the gas residing at the implantlocation or incorporate it in the implant volume. In either case, alarge gaseous void trapped by the implant in the nucleus of the disc istypically avoided for one or more of the following reasons: 1) itprovides a compressible volume into which the cured implant may shift orextrude, 2) the region in which the bubble resides is not load bearingresulting in localized forces on the vertebral endplates, 3) provides aspace where infiltrating cells are likely to accumulate and causeinflammation, 4) the gas entrapped in the space is likely to eventuallymigrate out of the disc space and cause a loss of disc height, and 5)the space creates a discontinuity between implant and annuluspredisposed to the creation of implant particulate. In one embodiment,the implant is space filling. In another embodiment, the implant issufficiently hydrophilic and/or foaming, wherein the final cured volumeis greater than the initial implant volume.

In one embodiment, the access (or opening) to the nuclear space of adisc is as small as possible. Often the therapy involves removal of someor the entire disc nucleus. The nucleus is removed and replaced by airduring the nuclectomy. The constraints of the access hole dimensions canmake it difficult to refill the nucleus with a liquid implant withouttrapping some gas in the nucleus. The region of trapped gas becomes avoid in the cured implant. This void can decouple a large fraction ofthe implant surface from the annulus making the implant less effectivein translating axial forces to radial forces applied to the annulus. Thebenefit of stiffening the annulus through the application of radialforces is incompletely achieved.

A large dimensional void in the implant, unlike uniformly distributedsmall bubbles, presents a high stress surface that can result infracture of the implant into the void and particulate formation. Thelarger the radius of the void, the higher the stress accumulation on thesurface of the implant and higher the likelihood the implant willdegrade. It is desirable to minimize voids in the implantation volume.

A non-expanding implant will push the air pocket to the side. Anexpanding implant will expand into the void. In the case where a volumeof gas is entrapped by the injection of a liquid implant, the initialpressure of the entrapped gas is at ambient pressure, and increases asadditional implant is injected into the space. During the injection, thepressure of the liquid implant equals the pressure of the gas trapped inthe void. For an expanding implant, the creation of a gas phase in theimplant makes the pressure inside the implant slightly higher than thepressure in the void. The implant foams into the void. While the mass ofthe trapped gas does not change, its distribution within the implantbecomes more uniform by becoming distributed in the bubble structure ofthe implant. This effect combined with the hydrophilic nature of theliquid implant tends to ensure the solid phase of the implant is incontact with the entire inner surface of the annulus and endplates.

In one embodiment, the implant fills the implantation space by acombination of increasing pressure after the implant volume has beendelivered and a space-filling foaming action.

Another embodiment provides a method for preventing migration of thefully cured implant. In most surgical applications, a hole will becreated surgically in the annulus of the disc through which degradednucleus is removed and implant is delivered. It may be beneficial toinclude at least a temporary tissue bonding means in the chemistry ofthe implant to prevent extrusion of the material through the hole in theannulus. In one embodiment, the implant does not fill the hole in theannulus. In one embodiment, this bonding can achieved by devices andmethods of Provisional Application No. 60/930,104, entitled “Foam DiscProsthesis” the disclosure of which is incorporated herein by reference.A tissue bonding implant of this embodiment can be useful in one or moreof the following ways: 1) bonds can be formed between the implant andinner wall of the annulus to stabilize and strengthen the annulus, 2)bonds can be formed between the implant and remaining nucleus localizethe nucleus and reduce the likelihood nuclear material will escape fromthe disc, 3) bonds formed between the vertebral endplates and layers ofthe annulus can allow for a lower modulus implant without extrusion, 4)these same bonds can maintain the patency of the surgically created holein the annulus allowing for natural annulus regeneration, and 5) theintegrated resulting structure of coupled layers of annulus, nucleus,and endplates can mitigate against the need for complete removal ofnuclear material and could potentially result in improved surgicaloutcome.

An in situ curing liquid implant that bonds to tissue renders it lesslikely to move around in the implantation space, especially if a voidexists in the implantation space. Motion of a nucleus implant within theannulus of a disc is generally undesirable because the differentialmotion between tissue and implant results in tissue disruption,inflammatory response and implant abrasion. A relatively low modulus,tissue-bound implant will follow the changes in tissue geometry andavoid particulate formation caused by tissue passing across the implantsurface. The ability of the implant to follow tissue motion requires animplant having a modulus similar to that of tissue; otherwise shearforces may develop between implant and tissue. In one embodiment wherethe modulus of the tissue matches the modulus of the implant, shearforce is minimized and the implant follows tissue motion.

Since the modulus of tissue is low, an implant of tissue-like compliancewill in many cases deform through a defect made in the annulus. Forexample, when forces are applied to the implant, and the modulus of theimplant is low enough to follow tissue motion, the implant will tend toextrude through a hole made in the annulus created during implantation.The existence of a defect in the annulus can re-establish shear stressbetween the implant and surrounding tissue, and the implant may moverelative to the tissue interface in a direction that favors extrusion ofthe implant from the implantation site. If the entire implant surface isfree to move in this way, there is very little restorative forcespresent to keep the implant in its intended site. However, if some orall of the surface is bonded to surrounding tissue, bulk slippage of theimplant relative to tissue is reduced or minimized. Although surfacestress develops at the bond interface, the force required to extrude theimplant through a defect in the annulus is greatly increased. Therefore,bonding allows for a softer, extrusion-free implant as illustrated inTable 1. In one embodiment, the implant bonds to tissue at least duringthe first weeks after implantation, when the defect in the annulus isopen.

In one embodiment, the bond strengths range from 4 lbs/in² to about 25lbs/in².

In another embodiment, delivery of the implant is performed through alumen. In one embodiment the lumen is of minimal cross section, butsufficiently large to deliver the composition (e.g., liquid nucleusimplant) or compressible preformed foam by conventional methods, e.g., acatheter or a syringe or similar liquid dispensing device that can beeither mechanically pressurized or manually pressurized sufficiently todeliver the liquid nucleus implant to the treatment site before theliquid implant has cured.

In one embodiment, the composition is a low viscosity liquid implant. Inone embodiment, the prepolymer is premixed with saline beforeimplantation to initiate the polymerization process and reduce theviscosity of the prepolymer. Reducing the viscosity of the prepolymercan have one or more of the following features: 1) a smaller deliverylumen and a less disruptive surgical procedure is possible, e.g., withrespect to preserving the structural integrity of the annulus, 2) theforce required to deliver the liquid implant to the implantation sitedecreases with decreasing viscosity, 3) the time required to deliver theliquid implant to the implantation site decreases with decreasingviscosity, and 4) a low viscosity implant ensures the crevices andfragments of the interior disc space are uniformly and completely coatedwith implant. Premixing the prepolymer with saline can have one or moreof the following features: 1) mixing with saline can ensure uniformactivation and curing of the implant, 2) mixing with saline can ensure amore predictable cure time, 3) mixing with saline can reduce the curetime, 5) mixing with saline can establish the liquid fraction of thecured implant, 6) mixing with saline can initiate the release of the gasphase of the implant, and 7) mixing with a quantifiable volumetric ratioof prepolymer and saline can determine the final solid/liquid/gas ratioof the implant.

In one embodiment, a relatively large implant volume of a liquid in situcuring implant can be inserted or delivered through a small hole.Delivery of the implant can be performed via injection, e.g., generallythrough a needle. The needle can be as long as several inches. In oneembodiment, the liquid implant is delivered under pressure; typicallypressures attainable by manual compression of a standard 3 ml syringeare acceptable, generally less than 100 psi. Low implant viscosity hasother benefits, for example, more of the injection pressure is availablefor pressurizing the implant to augment disc height. In some cases, itmay be beneficial to be able to sense resistance when injecting into thedisc, and the pressure increase due to resistance relative to the dropthrough the needle is related to the level of manual detection ofpressure resistance.

For the range of needle lengths and diameters commonly used to injectdyes into the disc for diagnostic purposes, an implant viscosity lessthan 1000 cp is acceptable, such as an implant viscosity of less than200 cp. In another embodiment, the viscosity ranges from 100 cp to 1000cp. In one embodiment, the viscosity limit of 1000 cp is satisfied whenprepolymer is mixed with water in the ratio of 70:30 or less, and 60:40or less for the 200 cp limit.

Another embodiment provides a cure time sufficiently short to ensure theliquid implant stays localized to the implantation site and sufficientlylong to allow the surgeon to deliver the implant in a beneficial way. Inone embodiment, the liquid implant cures at a faster rate when heatedinside the body and in the presence of tissue. In one embodiment, suchan implant can result in one or more of the following: 1) shear thinningof the liquid implant as it permeates defects in the annulus can causethe implant to warm via internal body temperatures and initiate proteinbinding at a faster rate than the bulk polymerization of the implant, 2)this polymerization mechanism tends to encircle the implant volume withcured implant to prevent loss of the implant from the nuclear space, and3) it can ensure that a large number of the active sites in the liquidimplant are available for covalent tissue bonding and not consumed inbulk polymerization. Implants without this feature that form solidimplants in situ, rely on mechanical attachment to tissue. In oneembodiment, a chemically enhanced bonded implant can tolerate greaterchanges in implant volume and shape before implant dislocation.

In one embodiment, cure times can span as long as 1 hour and as short as30 seconds, although longer cure times are possible. In general, accesswill have been made to the implantation site, and preparation of theimplantation site completed before the liquid implant is prepared. Inone embodiment, the implant is prepared by mixing between two syringesbridged by a female-to-female luer lok connection prepolymer in onesyringe and saline or other suitable aqueous solution in the othersyringe. In one embodiment, the hydrophilic nature of the prepolymerachieves homogenous mixing in approximately 10 mix cycles for mix ratiosof 10-90% prepolymer. In one embodiment, all implant ratios arehomogenous after 20 mix cycles. In one embodiment, the fastest cure timeare achieved where the mix ratio is approximately 1:1. However, the curetime does not differ by more than 100% for all mix ratios.

The surgeon typically requires a cure time long enough to mix and injectthe liquid implant and short enough to provide for in situ curing withina few minutes after implantation. In one embodiment, the cure timeranges from 1 to 10 minutes, such as from 3 to 5 minutes. In oneembodiment, the cure time halves for every 10 degree centigrade increasein mixture temperature. The typical difference between body temperatureand room temperature is about 10° C. Often, there is a decrease in curetime once the liquid implant is injected in the body.

Another embodiment provides an in situ cured implant that does notchange volume beyond a therapeutic range through the loss or acquisitionof aqueous fluid in the body. Many hydrogel compositions do not have across-linked structure, and when placed in an aqueous environment swellappreciably, sometimes to the point of dissolution. Polymeric swellwhere water enters the polymer matrix is distinguished from the phasetransition aspect of this embodiment where the implant increases volumethrough the liberation of a gas forming bubbles in the implant, whichmay later fill with water. The later aspect does not compromise thetensile strength and other mechanical characteristics of the solid phaseof the implant. Polymeric swell can compromise permanence of thepolymer, making them susceptible to fracture, migration, particulateformation, and loss of therapeutic efficacy. In one embodiment, the massof aqueous solution mixed with the prepolymer prior to implantation isapproximately the mass of water contained in the polymer matrix afterpolymerization such that the ratio of polymer to water in the solidphase of the hydrogel remains approximately constant.

In one embodiment, the foaming of an in-situ curable implant pressurizesthe implant volume and reduce voids in the implantation volume. Thefoaming may also aid in tissue infiltration and bonding. Moreover, thereare many preformed nucleus implants that expand after implantation. Thisfeature can minimize the access hole to the nuclear space by introducinga swellable preformed implant. The need for post-implantation swell isreduced for in situ cured implants since the implant volume can bepressurized prior to curing and the liquid nature of the implantprovides for small access holes. The primary disadvantage topost-implantation swelling of an implant is uncontrolled forces appliedto annulus and endplates that may result in annulus tears, endplatedeformation, implant extrusion or unnatural disc distraction.

In general, the chemistries can be adjusted to swell when placed in thebody. The tensile strength of a non-degrading hydrogel implant decreasesproportionally to the change in implant volume due to a decrease in thenumber of cross-links per unit volume. Swelled polymers can benoticeably easier to disaggregate, leading to an increase in particleformation. This aspect of a swelled implant may be disadvantageous whenall other endpoints such as disc height adjustment, implant sitefilling, implant modulus, as well as others are achieved in the curedstate of the implant.

In one embodiment, nucleus implant volume change affecting tensilestrength of the cured implant is less than 50% after 1 week for implantsplaced in the nuclear space of a disc. In one embodiment, the volumechange is less than 20% after 1 week, or less than 10% after 1 week.These limitations can relate to but are not limited to implant swell inthe body and increases of implant volume that results in decrease of amechanical property of the implant. This limitation may apply lessdirectly to implants designed to accommodate to the implant site withina predetermined therapeutic range, thereby providing tissues time toequilibrate with forces generated by the implant on tissue surfaces.Where implant volume decreases or increases in response to anaccommodative endpoint designed into the implant, that does not affectthe mechanical properties of the solid phase of the implant arecontemplated later in this application. These changes refer toalteration of the gas phase of the implant, and in certain changes inthe Poisson ratio of the implant after 1 week.

Another embodiment provides a nucleus pressurizing capability of theliquid nucleus implant. A liquid implant composition can liberate a gasphase during polymerization within the disc nucleus to facilitate itsspace filling aspects. In this embodiment, the choice of the ratio ofwater to prepolymer can determine the molar quantity of gas liberatedper unit volume of prepolymer and water. A therapeutically beneficialvolume of gas is liberated into the liquid prepolymer to cure into alarger volume foam implant to pressurize the nuclear space. Pressurizingthe nuclear space for in situ polymerizing implants can be achievedmechanically by applying pressure to the implant before it has cured insitu, and can be beneficial for one or more of the following reasons: 1)the pressure can be generated within the volume of the implant, notexternally, which enhances its therapeutic value by not disrupting aninterface formed between tissue and implant which may be beneficial tolocalizing the implant, 2) the pressure can be generated isotropicallywithout spatial gradients, 3) the pressure can directs the implant intointimate contact with the layers of the annulus, 4) the pressure can beeffective in achieving a higher implant modulus that is self-tailored tothe patient's condition and anatomy, and 5) the pressure can predisposethe vertebral bodies to distraction thereby achieving a greater discheight.

In one embodiment, the prepolymer compositions provide for gas to curedpolymer volume ratios of 3:1 greater under ambient conditions. In oneembodiment, the prepolymers possess an isocyanate functionality thatreacts with water to liberate carbon dioxide, resulting in a foam thatforms a pressurized cured implant state.

The pressure developed in the cured state can depend on how much theimplantation volume expands in response to the pressurized state of theimplant. In one embodiment, a beneficial outcome of the pressurizedstate is that the disc endplates become distracted and the height of thedisc increased. The pressurized state of the implant can increase themodulus of the implant, such that for a given volume it is possible tocalculate how much gas must be liberated so that the final gaseousfraction of the implant is at a certain target pressure. Thesecalculations involving molar quantities of gas liberating groups in theprepolymer are well known in the art.

Some of the advantages of the pressurized state of the implant may betransitory in nature, for example an initial pressurized state in thecuring implant that drives the isobaric expansion of the curing implantinto the small-scale features of the surrounding tissue, while the finalstate of the implant is insignificantly different from ambientconditions.

Another embodiment provides a composition further comprising aradio-opaque agent or illuminating marker. There are a large number ofradio-opaque and radio-emitting markers that are added to medicaldevices to aid in their external visualization by noninvasive means.They fall into two categories: soluble additive, and insolubleadditives. It is a requirement of either that they do not adverselyaffect the curing mechanism of the implant. It is further desirable thatthe additive not significantly adversely affects biocompatibility.

In one embodiment, the prepolymer is soluble in aqueous solutions, andthe radio-opaque agents is soluble in water. In one embodiment, theaddition of these agents to the aqueous phase of the mixture does notinterfere with implant polymerization. Optionally, a solid phaseradio-opaque agent may be added in the form of powder or aqueoussolution, which then forms a suspension upon mixing and does notinterfere with implant polymerization.

An example of a water-soluble radio-opaque agent is barium sulfate,which can be added to the liquid implant in volumetric fraction of 30%or less, or a volumetric fraction of 20% or less, or a volumetricfraction of 30% or less or 20% or less but not less than 10%. The bariumsulfate can be added to the aqueous fraction of the components to bemixed to form the liquid implant. The lower fractional proportion isapproximately the minimum amount of additive required to differentiatethe implant from surrounding bone and tissue, greater fractional amountsof additive increase the implant contrast.

An example of an insoluble radio-opaque agent is tantalum powder.Tantalum powder can be added to the aqueous or prepolymer fractions inpreparing the liquid implant. In one embodiment, the tantalum is presentin a volumetric fraction of 30% or less, or a volumetric fraction of 20%or less, or a volumetric fraction of 30% or less or 20% or less but notless than 10%. The tantalum powder can be of any size that is easilytransported through the delivery apparatus, e.g., sub micron in size.Although powder additions may present a potential for particulatemigration, a curing aspect or non-migration aspect of the implant canmitigate against particulate loss.

Another embodiment provides an implant with a bimodular compliancecharacteristic. The bubbles entrapped in the cured nucleus implant havea compliance described by the state equation for a gas. The complianceof the solid phase of the implant has a distinct and independentlyderived compliance that is governed primarily by the degree ofcross-linking in the formed implant and the amount of water comprisingthe hydrogel. When placed under load, the compliance aspects of each ofthese components sum in proportion to their volumetric ratio to yield abimodular compliance for the implant as a whole. One feature of thisbimodular aspect is that while the compliance for the solid phase of theimplant is expressed as a constant the compliance of the gas phase ofthe implant depends on the total volume of the gas phase. As the discheight decreases under load, the implant becomes stiffer—a feature notattained with implants with a Poisson ratio of approximately 0.50. Theimplant can freely deform and translate axial forces to radial forcesapplied to the annulus under small disc compression but can becomeincreasingly stiffer translating fractionally less of the axial forcesto radial forces under greater compression of the disc. This feature isprotective of the annulus, and guards against re-herniation of theannulus.

This bimodular compliance acts under small disc compression like thenatural nucleus of the disc, and acts increasingly more like a solidload-bearing member under increased compression of the disc. Thisimplant is engineered to be a compromise between distortion of theendplates typically experienced with harder nucleus implants, andherniation of the annulus or extrusion of the implant typicallyexperienced with the natural nucleus of the disc. Further, the bimodularaspect of the implant can be tailored by the medical professional to thespecific condition of a diseased disc by adjusting the ratio of aqueousand prepolymer fractions before mixing; or, the bimodular aspect can beengineered by the manufacturing concern to provide greater or lessermolar quantities of liberated gas during the polymerization of theimplant in situ.

In one embodiment, the implant (e.g., a preformed or cured foam implant)contains a solid hydrogel phase and a gaseous phase enclosed in thehydrogel in the form of bubbles. In one embodiment, the solid phasecomprises polyurethane and the gaseous phase is carbon dioxide.Depending on the method and amount of liquid implant delivered to theimplantation site, the final state pressure of the carbon dioxidebubbles enclosed in the closed cell structure of the hydrogel variesfrom near ambient to several atmospheres. The long-term pressure of thecarbon dioxide fraction of the implant depends on the solubility of thegas phase into the water residing in the hydrogel implant.

In one embodiment, the closed cell bubbles initially contain less than50% of a liquid, less than 25%, or less than 10% of a liquid (or mostlygas and minimal liquid) within, e.g., 1 month or less afterimplantation, within one week or less after implantation, within 5 daysafter implantation, within 3-5 days after implantation, within 3 daysafter implantation, or even within 1 day of implantation. In oneembodiment, the closed cell bubbles contain In one embodiment, at thesetime periods the Poisson ratio is less than 0.5. In another embodiment,after any of the above-mentioned time periods, the implant has a Poissonratio approaching 0.5.

At any instance in time, the compliance of the implant is a function ofthe combined effects of the compliance of the gas bubbles and theYoung's modulus of the hydrogel. Depending on the pressure in the gasphase of the implant, the dominant compliance can be approximately asthe ideal gas law for low gas phase pressure to the Young's modulus ofthe hydrogel for high gas phase pressure. In general, the compliance ofthe gas phase is softer than the compliance of the solid phase. Duringcompression of the implant, a point is arrive at where the pressure inthe bubbles exceeds the modulus of the solid phase, and the implant isessentially incompressible, and a Poisson ratio of 0.5 is approximated.

In one embodiment, the total compliance of the implant is bimodular,where the compliance is dominates by the gas phase for smallcompressions and is dominated by the solid phase for large compression.The early phase compression of the implant is characteristicallydependent upon the volume of the air bubbles, and thus increases withdisplacement. The late phase compression of the implant exhibits acharacteristic constant modulus, approximately by the Young's modulus ofthe hydrogel fraction.

The early phase compression can provide for high frequency loadabsorption without significant distortion of the annulus, or at leastless distortion than what would have resulted for a Poisson ratio=0.5implant. Therefore, one can image high frequency oscillations about amean disc height, and the mean disc height only changing when the meanload changes. This is protective of the annulus, and accommodative withregards the distribution of load between the endplates and the annulus.

Depending on the magnitude and duration of mean loads applied to theimplant, a volumetric fraction of the carbon dioxide in the bubbles canabsorb into the water fraction of the hydrogel matrix, thereby reducingthe pressure in the bubbles of the implant. This tends to restore someof the dynamic compressive range of the implant even when relativelylarge loads are endured for long periods of time. However, when theimplant is unloaded, the locally dissolved carbon dioxide can reappearin gaseous form and depressurizing the bubbles of the hydrogel. However,this response is dissipative with time, and eventually the carbondioxide is lost by diffusion and the time averaged volume of the bubblesis filled with free aqueous solution derived from the surroundingtissue.

The long-term compliance of the implant is governed by the Young'smodulus of the hydrogel fraction, and eventually the implant exhibits aPoisson ration of 0.5. At this point height changes of the disc can bemodulated by height and radius changes of the annulus, and the implantadopts the compliance properties of the natural nucleus.

The short-term bimodular characteristic of the implant can result inisolating high frequency displacements of the vertebral endplates fromdisplacements of the annulus, the effect of which is to provide forannulus healing, especially with respect to healing of the access holecreated in the annulus for removal of nucleus and injection of liquidimplant. The short-term bimodular characteristic can establish theheight of the disc with respect to the partition of axial and radialforces, which is essentially locked into place when the bubbles arefilled with water and the compliance of the implant, approximates aconstant value. The disc height is not necessarily coupled to thecompliance of the implant before this accommodation can be achieved. Inone embodiment, this provides a method for axial forces to be somewhatdecoupled from radial forces until a mean disc height is established,such that the range of accommodation of the annulus can be centered onthe range of displacements created by loading whereby the radially andaxial forces are partitioned by the Young's modulus of the implant whenthe bubbles become filled with a liquid fraction derived fromsurrounding tissue

Another embodiment provides a liquid nucleus implant, the compliancecharacter of which is predictably controlled by the medicalprofessional. A medical professional may decide to surgically modify avertebral disc to correct a pathologic condition. Typical interventionsinclude: 1) reduction or removal of a portion of the disc annulus, 2)removal of a portion or all of a disc nucleus, 3) removal of a portionof the inner layers of the disc annulus, 4) modification of the distancebetween endplates of a disc, and 5) various procedures intended toassess the compliance or spatial aspects of the diseased disc. Theseinterventions can yield characteristics of the disc that can be used toassign the appropriate implant therapy. For example, the size of theannulus hole, the thickness of the annulus, the radius of the spacecreated in the nucleus of the disc, the height of the disc, the degreeof compliance of the diseased annulus, and the dimensions of theendplates are all useful inputs into a therapy paradigm. The equationsinvolving these inputs can be used to select a set of ideal implantcharacteristics as set forth in the Appendix below. These equationsdemonstrate that the type and range of implant characteristicsselectable by the medical professional is adequate to address a ratherbroad range of pathological disc conditions.

In one embodiment, the prepolymers of the composition are hydrophilicand readily dissolve in water. The lower limit for solid formation isapproximately 5% by volume prepolymer in a mixture substantiallycomprise of water. The limit on Young's modulus for the in situ curingimplant is the solid/liquid phase transition, which occurs in thepolymerization at the 5% concentration. The upper limit can depend onthe molecular weight of the prepolymer and/or its functionality, buttypical prepolymers reach maximum modulus at nearly 100% prepolymermixtures with only trace amounts of water. The water typically presentin the body can be more than adequate to complete polymerization of theprepolymer. In one embodiment, the upper limit on the modulus is in therange of 10-20 MPa, e.g., from 0.5 to 20 MPa or from 0.5 to 10 MPa. Inanother embodiment, the upper limit on prepolymer-water mixture able tobe delivered through a 20G needle with hand pressure is approximately 5MPa, with mixtures ranging from 0.5 to 5 MPa, 1 to 5 MPa, or 1-3 MParesulting from mixtures in the 10-80% water range. These ranges andtheir associated procured mixtures can depend on the cross-link densityachieved with the prepolymer structure.

Prepolymers, as those disclosed herein, are adequate to meet the rangeof therapeutic implant moduli detailed in Table 1. Additionally, fillerssuch as flock or particulate added to the procured mixture will tend toshift the range of moduli to higher values.

Since the modulus varies as the cross-link density it also variesapproximately uniformly with the mixture ratio. Accordingly, for anyparticular prepolymer a table can be generate for cured implant modulusvs. mixture ratio. This feature is useful in practicing the type oftreatment guidance illustrated in Table 1.

Another embodiment provides a liquid nucleus implant that isself-sealing. In this embodiment, the implant polymerizes preferentiallyalong the implant margins to achieve one or more of the following: 1)prevent extrusion of the implant during pressurization through theaccess hole made in the annulus, 2) prevent externalization of theimplant through undetected defects in the annulus, and 3) provide forpressure to develop within the volume of the in situ curing implant. Oneor more of the following clinical benefits can be achieved: 1) thelimited infiltration of the liquid implant into and along the concentriclayers of the annulus to enhance the hoop stress capacity of the annulusand enhance the localization of the implant, 2) to provide anencapsulating aspect to the outer layers of the implant that act torestrain its deformation into the access hole in the annulus and therebyleave it patent to promote annulus healing, and 3) to provide a layeredaspect to the nucleus implant that acts as a structural extension of theannulus and provides additional hoop stress bearing properties.

The rate of polymerization or curing of the prepolymers described hereincan vary according to two environmental conditions. In certainembodiments involving in situ curable implants, the cure rate increasesapproximately 2 fold for every 10° C. rise in liquid implanttemperature. This can be a useful feature since liquid implant injectedat room temperature will cure preferentially at the margins where theimplant is in contact with tissue and locally heated by the tissue. Thiseffect can be further magnified when the implant extends beyond the bulkof the implanted mass, and begins infiltrating cracks and fissures inthe tissue interface. The effect can be large enough to cause nearlyinstantaneous polymerization of the liquid implant as it infiltratesthis tissue. In cadaver studies it was observed that in discs withpartially delaminated annulus tissue, that the liquid implant followedthese fault lines and acted to recouple the layers mechanically. Thisdevelopment has two effects: 1) the maximum hoop stress of the annulusis dramatically increased since the layers of the annulus are no longerfree to slide past one another and concentrate stresses, and 2) theinward projection of the inner layers of the annulus, creating aconvexo-convex cross section is defeated and the normal convexo-concavecross section is restored due to the radially directed pressuredeveloped in the curing implant.

The second environmental condition that contributes to the self-sealingfeature derives from the difference in reaction rates of polymerization.For example, where the prepolymer comprises isocyanate groups, theself-sealing feature derives from the difference in reaction ratesbetween the reaction of the isocyanate groups on the prepolymer withwater to form amine groups and the reaction rate between amine groupswith isocyanate groups to form urea linkages. The latter occurs muchfaster than the former. Since the cure rate of the liquid implant islimited by the rate of amine formation in the implant, amines existingon the surface of tissue will react first with the implant. Thus, thecure rate at the periphery of the implant can be increased by the actionof the water component of the liquid implant that generates amine groupsand the preexisting proteinaceous groups residing either freely orattached to the tissue. This acceleration of the cure rate can befurther increased when the implant infiltrates tissue porosities, sincethe ratio of implant surface area to volume is increased. In the limitthe rapidly polymerizing layers of the implant come together in thenarrow confines of a tissue defect causing the implant to bond tissuelayers together and seal against the egress of liquid implant.

This self-sealing feature can be useful in the clinical setting wheredefects in the annulus may go undetected, and externalization of theimplant outside the disc would be deleterious to outcome. In oneembodiment, the sealing aspect is achieved prior to substantialliberation of gas within the implant, allowing the seal to be achievedbefore full pressurization of the implant.

Another embodiment provides a nucleus implant that forms a foamedhydrogel in the body capable of exchanging water occupying the solidphase of the implant with surrounding tissue. The water is weakly heldwithin the polymer matrix, and can be exchanged with water molecules ofthe body. In one embodiment, the exchange can be one-to-one, such thatthe total volume of the solid phase of the implant does not changesignificantly. The density of the polymer matrix responsible forlocalizing the water is determined largely at the stage of mixing, andis determined more rigorously at the completion of polymerization.Although most tissues present a wet aspect in the body, the free wateron these tissue surfaces is generally insufficient to altersignificantly the ratio of water to prepolymer during the polymerizationphase of the implantation. The water exchanging aspect can provide oneor more of the following clinical benefits: 1) ionic or pH differencesbetween the implant and the surrounding tissue quickly equilibrate, 2)osmotic differences between the implant and surrounding tissue quicklyequilibrate, 3) accordingly, nutrient conduction to distant tissues fromsources in contact with the implant, such as the vertebral endplates, isnot blocked by the implant, and 4) proteins and cells responsible formarking and proliferating an inflammatory or fibrotic response arehindered by the surface washing aspect of the implant resulting from theconstant thermodynamic exchange of water molecules on the implantsurface.

In one embodiment, the cross-link density of an in-situ curable implantis in part determined by the mixture ratio of water to prepolymer. Inanother embodiment, the prepolymer is hydrophilic and all the water isincorporated into the cured implant for prepolymer ratios greater thanapproximately 5% water. The mode of incorporation is via loose hydrogenbonds that are reversible under statistical fluctuations and heating.However, the vacancy created by thermal excitation of a water moleculeresulting in its diffusion out of the implant is preserved in the curedimplant structure, making it thermodynamically more probable that thevacancy will be filled by a water molecule originating from thesurrounding tissue. This feature is one reason for the negligible swellthat occurs when the cured implant is placed in a bath of water and itsvolume stability within the body.

This constant exchange of water molecules with the environment providesfor nutrients, ions and water to pass through the cured implant,principally by diffusion. This feature has the benefits of equilibratingthe chemistry of the implant with the chemistry of the surroundingtissue, allowing nutrients to pass from tissue layer to tissue layeracross the implant, and discourages attachment of proteins that markimplants for foreign body response. This feature can render the implantuniquely biocompatible, and can reduce a significant foreign bodyresponse without interfering with transport mechanism naturallyoccurring in the disc.

Accordingly, another embodiment provides a nucleus implant that forms afoamed hydrogel in the body capable of allowing complex moleculediffusion through the implant volume. This aspect can be beneficial intwo ways: 1) it permits the natural flow of ionic and molecularconstituents between tissues, and 2) provides for diffusion out of theimplant therapeutically useful chemical structures. These structures canbe distributed uniformly in the implant during mixing. These structurescan be water soluble or suspended. The diffusion of these structures outof the implant is controlled by diffusion, and the rate of diffusion canbe controlled by adjustment of certain characteristics of the implant.In one embodiment, the implant can act as a drug delivery device, sourceof cell directing proteins such as growth hormones, or an initiator ofcellular responses. In another embodiment, the implant can re-hydratethe often dehydrated condition of a diseased disc.

Another embodiment provides a permanent (i.e., nonabsorbable) implant.The permanence can be superior compared with protein-based and manypolyol-based in situ polymerizing implants. Implant permanence can beachieved, in one embodiment, by decreasing one or more of the magnitudeof implant swell, hydrolysis, oxidation, and chain breakage.

In one embodiment, the implant is an in-situ curable compositioncomprising prepolymers that form polyurethane within the body andpossesses permanence characteristics similar to those of polyurethanes.There are two classes of polyurethanes used commercially: polyesterurethanes and polyether urethanes. In one embodiment, the cured implantscomprise polyether urethanes. Polyester urethanes are known to degrademore rapidly in the body. Even among polyether urethanes, there areseveral chemical variations that contribute to implant permanence. Manymedical grade polyurethanes are not cross-linked so they can bedissolved in solvents and applied to medical devices as coatings.Polyurethanes that are not cross-linked swell in the body and degradefaster. In one example, the prepolymer has a multifunctional structurethat results in a crosslinked implant. In this example, the structure ofthe prepolymer comprises a small multifunctional center joined tohydrophilic arms through hydrophobic linkages. One or more of the ratioof hydrophobic to hydrophilic groups and their distribution within theprepolymer, the choice of functional end groups, the molecular weight ofthe prepolymer, the absence of available OH groups in the prepolymer andthe fully functionalized condition of the pendant structures of theprepolymer, the structure of the bonds formed during bulk polymerizationof the implant, and the presence of a biologically insignificantquantity of free functional groups, can be tailored to achieve apermanent implant.

During in situ polymerization, the chain extends to prepolymer and alsoforms the cross links resulting in a molecular structure that ismillions of Daltons in weight. Theoretically, in situ polymerizedpolyurethane can form a single polymeric molecule. Highly cross-linkedsingle macromolecules do not dissolve in solvents and do not melt withheat. This resistance to phase change results in their improvedpermanence in the body.

Among the highly cross-linked polyurethanes, there are certain typesthat are more stable in the body. In one embodiment, the use of anaromatic isocyanate rather than an aliphatic isocyanate can improveimplant permanence. In one embodiment, the prepolymers possess one ormore of the following structures that contribute to implantpermanence: 1) the prepolymer is functionalized with aromaticisocyanate, 2) the backbone of the prepolymer contains both polyethyleneand polypropylene units, 3) the prepolymer is multi-functional, and 4)the multifunctionality is achieved by isocyanate chain extension with alow molecular weight triol.

Another embodiment provides a single component, in situ curable liquidimplant, i.e., contains only one functional or reactive component.Currently, there are a number of in situ polymerizing implants used inthe disc and other areas of the body that involve combinations of two ormore reactive components that are necessary in the polymerization of theimplant. In situ forming implants that rely on the union of two or morefunctional groups in order to form the implant, risk certainbiocompatibility features, including loss of some fraction of each ofthese components to surrounding tissue. When these components becomeseparated in the body, their characteristic chemical activity canpresent a biocompatibility risk. Frequently the separated components acton proteins within the body that result in immunotoxicity,sensitization, cellular toxicity, carcinogenicity, teratogenicity, andmutagenicity. These biological responses can contribute to adverseevents in the clinical population.

In an embodiment that provides a single component in situ curableimplant, if the prepolymer migrates from the implantation site, it canrapidly polymerize with itself into a biocompatible configuration. Itsbiocompatibility is not dependent upon the reaction of two or moredistinct components that may become separated and migrate systemicallywithin the patient. In one embodiment, the single component compositionpolymerizes in situ without the addition of water or any other reactivecomponent. Any water addition is made to achieve aspects not central tothe in situ curing feature; e.g., achieving a suitably low viscosityimplant, achieving a desired modulus for the cured implant, and otheraspects.

Single component in situ curing materials rely on some aspect of theimplant preparation or implantation site chemistry to achieve curing,since their one-part formulation must be stable in its packaging. Theadvantages of single component implants is that the active structures inthe pre-cured implant do not rely on the coupling together of two ormore active structures to form a biocompatible implant. They achieve acured implant by a transformation of a fraction of the implant volumeinto an active form. As each molecule is transformed it is rapidlyconsumed by the remaining un-transformed species. In one embodiment, theprepolymer cures first at the tissue interface, resulting in the uncuredimplant to be quickly isolated from the body before the ratio ofactivated structures reaches parity with the un-activated structures.Parity between activated and un-activated structures presents theopportunity for these parts to become isolated from each other withinthe body.

In one embodiment, the activated structure of a single componentcomposition reacts more rapidly with inactivated structures than therate at which the implant environment or implant preparation createsactivated structure, resulting in more inactivated than activatedspecies, at least in the beginning stage of polymerization. In anotherembodiment, the formation of biocompatible implant occurs first at theperiphery of the implant volume so that unreacted species are contained.

Another embodiment provides an in situ polymerizing compositioncontaining a high molecular weight polyethylene glycol. In most in situpolymerizing structures, including those disclosed herein, active groupsare employed in the polymerization of the implant that are alsopotentially reactive with chemical, and cellular structures in the body.Attaching these active groups, which are typically low molecular weight,to the high molecular weight PEG can realize one or more of thefollowing biocompatibility benefits: 1) larger molecules diffuse moreslowly through tissue and liquids and thus are less likely to circulatesystemically in the body, 2) the use of larger molecules reduces themolar fraction of active groups in a composition, and 3) thebiologically active aspects of a functional group are often reduced byseveral order of magnitude when coupled with a large, biocompatiblemolecule. In one embodiment, the high molecular weight polyol has amolecular weight of several hundred or more Daltons (e.g., 200 D ormore, 300 D or more, or 500 D or more). Reductions in toxicity asmeasured by LD50 assessments can be as large as an increase by 10,000the mass of the compound required to achieve an LD50 level of toxicity.

In another embodiment, the prepolymer molecule, or collection ofmolecular structures possess a high molecular weight so that they do notrapidly diffuse into tissue and that their toxicity is reduced. It iswell known that attaching polyethers to low molecular weightdiisocyanates reduces their tissue toxicity, often by several orders ofmagnitude. The prepolymers, such as those of Example 3, can be formed byfunctionalizing diols of approximately 1000 Dalton with aromaticdiisocyanate and chain extending with a triol to form prepolymer specieswith a mean molecular weight of 4,000 to 6,000 Dalton.

Another embodiment provides a prepolymer of sufficient hydrophilicity.For example, prepolymers comprising polyols functionalized withisocyanates are typically hydrophobic, and capable of retaining lessthan 10% by volume water in the polymerized structure. In embodimentswhere the prepolymer comprises isocyanate groups, depending on thechoice of isocyanate group, the prepolymer is capable of forming ahydrogel comprising at least 70% water by volume. This feature canprovide one or more of the following benefits, including: the strengthof tissue bonding, the affinity for filling spatially small features ofthe implantation space and filling the space without voids, as well asenabling the ratio of water to prepolymer in the liquid implant to beadjusted by a medical professional to achieve a desired treatmentoutcome. The prepolymer component can be attracted to water aggressivelyand can be drawn into tissue features by the presence of water, e.g.,when the mixed fraction of water is small or the prepolymer is useddirectly. When prepolymer is used alone as a pretreatment to a laterimplantation of mixed prepolymer, the prepolymer can rapidly coats andseals the tissue surfaces of the implant location. The presence ofproteins on the tissue surface and the mixing that occurs as theprepolymer is drawn across the tissue surface can serve to acceleratepolymerization and the formation of a sealing surface. When used in thisway sparingly with water, the resulting modulus of the formed layers canbe more rigid than an implant formed by mixing with a more generousquantity of water. In this two-step application of the implant, acompound structure can be formed in situ comprising a rigid outer shelland a soft inner core.

In one embodiment, the prepolymer is a polyurethane prepolymer. Suchprepolymers do not generally form hydrogels with water content exceedinga few percent by volume. Many do not polymerize in the presence of waterat all, but those that do generally form rigid solids. For example,moisture curing polyurethane varnishes form durable coatings with highdurometer. In one embodiment, the prepolymers form a foaming hydrogelwithout employing hydrophobic components. In another embodiment, theprepolymers form foams that do not possess tensile strength suitable forindustrial foams, coatings and the like. In another embodiment, the lowmolecular weight diisocyanate is consumed in end capping the polyetherstructures. In yet another embodiment, polyurethane prepolymer solutionscontain between 5 and 25% or more of free, low molecular weightdiisocyanate. Such a composition can form soft, pliant yet chemicallyresistant fully cross-linked polymers. The prepolymers can minimizebiological toxicity by minimizing the free isocyanate content andbinding active isocyanate groups to high molecular weight chains. Inanother embodiment, polyurethane prepolymer solutions contain between 1and 5% of free, low molecular weight diisocyanate (e.g., less than 500D, or less than 200 D).

Isocyanates are hydrophobic as a class, and the grafting on of highmolecular weight polyethers not only increases their biocompatibilitybut also increases their hydrophilicity. This feature can achieve ahomogeneous mixture of prepolymer and water in ratios significantlyabove a few percent. For example, combining isophorone diisocyanate witha 4500 Dalton triol composed of units of polypropylene oxide andpolyethylene oxide in the ratio of 1:1 results in a prepolymer thatprecipitates out of solution before polymerizing at mix ratios above 50%water. However, combining toluene diisocyanate with a 4500 Dalton triolof pure ethylene glycol dissolves readily in any ratio with water, butthe polymerization product swells greater than 100% by volume in water,exceeding this value for high water mix ratios. Thus, there is acompeting condition between the potential for forming a high watercontent hydrogel and forming a volume stable implant.

Prepolymers formed using polyethers with polyethylene and polypropylenein the ratio 1:1 do not generally wet tissue well, and may not makecontact with tissue sufficiently to provide for the formation of goodtissue bonds. Furthermore, hydrophobic prepolymers require longer curetimes; many compositions have cure times outside the 2-10 minute rangerequired in medical applications. In balancing the hydrophilicity of theprepolymer, the characteristics of the polymeric backbone and theisocyanate can be considered together. For example, in building largeprepolymer molecules to achieve an acceptable level of biocompatibility,in some embodiments the compositions intersperse the molecule with hardsegments of the type obtained by chain extending two polyols with adiisocyanate. The existence of these hard segments can contribute toimplant permanence and volume stability. In one embodiment, thecomposition comprises diols containing polyethylene and polypropylene inthe ratio 75:25, respectively and chain extend them using toluenediisocyanate and trimethylol propane to achieve a trifunctionalprepolymer with nitrogen contained centrally in the prepolymer moleculeand a mean molecular weight of approximately 5000 Dalton (e.g., between4000 and 6000 D, e.g., between 4500 and 5500 D).

Another embodiment provides an implant that can bind the cured foamspreviously formed by a first application of the prepolymer composition.Accordingly, this embodiment is directed to multiple applications (twoor more) of the liquid implant to an implant site. The multiple implantmaterials resulting from the multiple applications can be bondedtogether. This can establish a proper shape and volume of an implant toachieve a desired clinical outcome; for example, an increase in discheight. A clinician may revise an initial implant by adding to anexisting implant, either peri-operatively or at a future date. Suchrevisions may include treatment paradigms intended to maintain animplant's clinical efficacy, or to gradually direct a remodeling of animplant locus. For example, in the latter, staged additions of implantvolume could be administered in order to increase disc height withoutcausing failure in the surrounding tissue. The implant may be usedwithout causing annular herniation, annular tearing, cracking ordeformation of the endplates. The implant may be used protectively tosegregate one tissue type from another. For example, the implant may beused to prevent extrusion of natural nucleus from the nuclear space ofthe disc. The implant may be used to reinforce an annulus of a disc,wherein a first application is directed between the layers of a discannulus and a second application is directed into the nuclear space,with or without removing nuclear material.

The surgical applications can comprise one or more of 3 primary uses: 1)replacement of some or the entire disc nucleus, 2) repair and strengthenthe disc annulus, and 3) bond or fit a preformed nucleus implant to asurgically prepared site. In one embodiment, tissue bonds take twoforms: 1) mechanical bonds where the implant is anchored to tissue byinfiltrating small-scale tissue structures and solidifying within thesestructures, and 2) chemical bonds which include covalent bonds as wellas hydrophobic association and charge mitigated association. Theintegrity of these bonds depends on the stability of the implant.

In many applications, where the hole made in the annulus to removenucleus is relatively small, e.g., <4 mm, the bond strength of theimplant need not exceed 2-4 lb/in² in shear. However, the bond strengthrequirement has no upper limit and compositions that form stronger bondsare useful in a broader range of implant scenarios. In one embodiment, atissue bonding implant has a bond strength that equals or exceeds thetensile strength of the implant. Suitable bond strength in shear is inthe range of 10 lb/in² or higher, e.g., 25 lb/in² or higher.

Bonds strength and shelf life of the prepolymer can be enhanced byensuring all the hydroxyl groups are terminated with isocyanate groups.In one embodiment, an excess of low molecular weight isocyanate isprovided to ensure that all hydroxyl units are capped. The amount ofthese low molecular weight isocyanates in the prepolymer is a smallmolar fraction, ranging from 1% to 5% of the composition, and typicallyless than 1%. Larger fractions of low molecular weight may decrease theshort-term biocompatibility of the implant, although the amines of theselow molecular weight species are quickly eliminated from the body ifthey should escape the implantation site. In one embodiment, bondstrength is generally increased up to a limit of a few percent of lowmolecular weight isocyanate.

Determining the precise amount of liquid implant to deliver to therepair site can be difficult to determine, and often relies on sensingback pressure in the syringe, observing spatial changes in the disc, orusing noninvasive imaging means. Even with perfect determination of atherapeutic volume, the implant and/or surrounding tissue is likely tochange in response to loading. In terms of long-term efficacy, tissue orimplant may degrade requiring additional implant volume to maintainefficacy. The minimally invasive nature of implantation of liquid insitu curing implants affords a possibly of revision. A change in themechanical properties of the spine may require implanting additionalmaterial over pre-existing material. To that end, multiple applicationsmay be needed where the subsequently applied liquid implant bonds to itscured form. In one embodiment, the bond between an earlier implantationand a later implantation is seamless, such that when the combinedimplants are tested to failure that the failure locus is other than thejoint between the two implants. In one embodiment, a textured surfacepresents an opportunity for mechanical bonding to a pre-existingimplant. In one embodiment, the prepolymer bonds both mechanically andchemically to its cured form.

Another embodiment provides one-piece removal of an implant formed insitu. The cohesive strength of the implant after polymerization can besufficient to allow clinicians to retrieve substantial portions, if notall, of an in situ formed implant in one piece.

In one embodiment, the bulk strength of the implant allows for cleanremoval of the implant if revision is required. Generally, implantssufficient to withstand the forces occurring in the nuclear space willsatisfy this objective. Implants that do not satisfy this requirementare those that degrade in the body so that it is difficult to determinethe boundary between implant and tissue. In general, the more permanentan implant is the easier it will be to remove. Other considerationsinvolve the biocompatibility of the implant. An implant that induces astrong or chronic inflammatory response may be deeply embedded infibrotic tissue such that removal of the implant risks damage tosurrounding tissue, for example nerves. Other aspects of an implant thatmake it difficult to remove are its propensity for migrating in thenuclear space. Implants that tunnel into the annulus or endplates makeremoval of the implant difficult, and risk leaving behind a far moreincompetent anatomy. In one embodiment, the implant features hightensile strength; biocompatibility, permanence, and low modulus toachieve an implant that can be remove years after implantation.

Another embodiment combines the features of tissue bonding,self-bonding, and affinity for water to achieve a liquid nucleus implantthat can be employed in the reconstruction of the annulus. For example,a thin annulus may be reinforced by coating the inner layers of theannulus, coupling the inner layers of annulus to a nucleus implant,mechanically displacing an annular surface and bonding to an nucleusimplant so as to correct a herniation, bonding together layers of anannulus to increase the annulus hoop strength or prevent extrusion ofnucleus between layers of an annulus. Lastly, the liquid implant of thecan be used to fill a void or defect in an annulus, correct a fissure orcrack, redistribute loading forces in an annulus.

In diseased discs it is sometimes the case that the transition from theless structured nucleus to the laminated structure of the annulusbecomes less distinct. The increase in disorder of the disc as a wholeleads to failure of the annulus. When the disc height is in the normalrange the orientation of the fibers in the layers of the annulusalternate in criss cross fashion from layer to layer. The angle of thefibers between two adjacent layers are closer to 90 degrees than 0degrees. In the diseased case, the height of the disc can decrease andthe angle between the fibers in adjacent layers of the annulusapproaches zero. When this occurs, the organized structure of theannulus can be compromised by the radial forces generated by axialloading of the disc, which in the collapsed state can result in fibersof one layer passing through fibers of another layer. When theorientation of fibers in adjacent layers becomes more parallel, thosefibers can migrate across layers resulting in herniation.

Whatever fraction of the radial force not contained by the first layeris passed on to successive layers. There is an advantage instrengthening the inner layers of the annulus because the inner layersare frequently the most disorganized. The containing capacity of theinner layers can be enhanced by erecting a barrier that resists fluidflow between the fibers of the annulus. An in situ bonding polymerapplied to the inner surface of the annulus can both diminish thelikelihood of fluid extrusion through the annulus, but also increase themaximum hoop stress of the layers. If the layers have already becomedisassociated, that is the layers can easily be dissection apart, thelocal expansion of the liquid implant limit infiltration along theselayers of least resistance. Cadaver studies have shown that smallamounts of liquid implant can flow between disassociated layers of anannulus and bond them together. If the disc height is restored duringthis bonding process, a preferred orientation of the fibers betweenlayers is achieved.

When a disc becomes compressed and the volume of nucleus in the discdecreases the outer layers of the annulus bow away from the disc center,and the inner layers can bow toward the disc center. This conditionresults in a force within the annulus that tends to separate layers ofthe annulus. Nucleus can then enter a defect in an inner layer; travelcircumferentially until it reaches another defect in the next layer andprogress in this way until nuclear material escapes the disc. Nuclearmaterial outside the disc results in a strong fibrotic response in anattempt to heal an abnormal condition of the spine. The fibroticresponse is rather ineffective in containing the nuclear material andcontributes to irritation of nerves surrounding the disc. In the extremecase, this growth of fibrotic tissue can affect the spinal cord. Theresult is a fissure in the annulus, which may be associated withmultiple branching, circumferentially directed paths leading out of thedisc. Removing nuclear material from the defect and filling the defectwith a tissue bonding implant can repair this condition. Not only isthis pathway blocked to further loss of nuclear material, but replacingthe shrunken nuclear material with synthetic nuclear material preventsthe inward bowing of the annulus and restores the cross sectionalgeometry of the annulus to a more normal shape.

Another embodiment eliminates the need for a molding element or balloonto isolate from tissue a toxic element of an in situ forming implant.Enclosing an in situ curing material within a distensible container sucha balloon isolates tissue from the curing reaction. While this approachis attractive from a safety point-of-view it may defeat many of thetherapeutic properties of such devices. Enclosing the liquid implant ina forming element may reduce, if not eliminate, the formation ofmechanical and chemical bonds between cured implant and surroundingtissue. This observation may apply also to any advantage derived fromstrengthening the annulus. Such enclosures are by their natureimpermeable, and thus are likely to block fluid flows within the disc.Lastly, motion of the implant within the disc can result in sheardegradation of the implant, particulate formation, and fibrosis andeventually externalization of the implant.

Another embodiment provides the absence of neurotoxicity associated withboth the prepolymer liquid implant and the polymerized result. Implantinjected into the nucleus in a liquid state presents the chance thatimplant escapes the nucleus through a defect in the annulus, and curesoutside the disc, possibly on the surface of surrounding nerves. Becausethe implants are chemically active, they could chemically interact withthe nerves compromising their function. In studies of rats, prepolymerwas placed on the exposed surface of rat brains and allowed topolymerize. The animals were monitored by EEG and behaviorally for 7days, and histopathology was obtained. Accordingly, in one embodiment,the prepolymers do not compromise nervous function.

Another embodiment minimal to no migration of the implant. Migration ofimplants and their derived particulates requires an active interactionbetween implant and tissue that results in implant displacement throughtissue. Factors contributing to implant migration include one or moreof: 1) the difference in modulus of the implant and surrounding tissue,2) the degree of inflammatory response initiated by the implant that canresult in tissue death and faults along layers of tissue, 3)differential motion between the implant and surrounding tissue, 4) thepropensity of the implant to chemically decompose or form particulates,5) lack of homogeneity of the implant, and 6) the propensity of theimplant to localize stresses on the implant surface.

The formation of bubbles in the cured implant can mitigate againstimplant migration. In one embodiment, the foam structure is closed cell,and occupies between 30 and 50% of the cured implant volume. In oneembodiment, prepolymers of approximately 5000 Daltons mean molecularweight possessing three functional NCO groups per molecule andpossessing a backbone comprising polyethylene and polypropylene in theratio 75:25 generate adequate anti-migration characteristics when mixedwith an equal volume of aqueous solution prior to implantation.

Another embodiment relates to the insensitivity of the therapeuticbenefit to implantation volume, implantation technique and implantpreparation. The formed nucleus implant can have a level-seeking featurethat tends toward a physiologically beneficial disc height. In practice,surgical improvements in back pain can be achieved by reversing thenormal reduction in disc height that occurs with age. Therefore, thetherapeutic bias is towards increasing disc height. Increasing theheight of a vertebral disc can be achieved by, first reducing the loadapplied to the disc annulus and second mechanically distracting thespace between adjacent vertebral bodies. Both of these steps riskdeforming the endplates of the vertebral bodies or herniation of thedisc annulus. Removing load from the annulus involves shifting loadtoward the center of the vertebral endplates. The endplates are concave,and forces applied closer to their center reduce the load bearingcapacity of the endplates. The endplates can fail suddenly or thenatural concavity of the endplates can be accentuated, resulting in theperipheral surfaces of the endplates moving together. There are otherstructures coupled to the endplates that also can cause back pain, suchas a change in the spatial relation of the facet joints.

Mechanically distracting the space between endplates is generallyachieved by applying pressure to the endplates, and thus this step alsorisks damaging the endplates. The damage is more frequently sudden innature, and may involve fracturing the endplates. Recently, strategieshave been contemplated where an implant is placed in the nuclear spaceunder pressure, or the implant swells to develop a pressure. In bothcases the goal is distraction of the endplates. In some cases the forcesare axially directed, but in other cases, especially where a balloon isused or a homogenous implant is formed in situ the forces tend to bemore isotropic, applying forces equally to the endplates and theannulus. This can contribute to herniation of the annulus resulting inthe painful condition of pressure being applied to nerves surroundingthe disc. These risks can be categorized as acute or chronic. Theirassociated adverse events are linked to non-accommodative features ofthe therapy that generally place non-physiological forces on the spineor create non-physiologic spatial relations between spinal elements. Theacute risks and associated adverse events can be avoided by applying thereconstructive therapy over a longer interval of time. Therefore, thereis a need for a means to apply corrective forces to the spine overextended periods of time. However, coupled with this need is the need toguard against applying forces that are excessive, and which may createspatial relationships between the various elements of the spine thatcause pain. These needs can be satisfied in an implant that is bothaccommodative and applies its beneficial aspects over time. An implantis accommodative if its mechanical characteristics are in partdetermined by the mechanical characteristics of the spine. In thisrespect, the bimodular aspect can be useful in applying axial loadingforces only when the disc height has deteriorate to a certain degree.

Another accommodative feature is the presence of bubbles in the implant.In addition to having a nonlinear compliance, they also have thecapacity to decrease in size depending on the loads applied to theimplant. For example, sustained loads tend to force the gas in thebubbles to dissolve into the liquid of the solid hydrogel component ofthe implant. Accordingly, the bubbles collapse with no additional loadneeded. A factor on whether the bubbles collapse is the duration of theapplied load. Conversely, when the load is reduced the gas dissolved inthe implant can reform as gas, re-establishing the additional complianceconstraint associated with re-inflating the bubbles in the implant. Thebubbles provide additional accommodative features. For example, overseveral days the implant is likely to relax to a certain disc height.The degree of relaxation will depend on the ratio of the components inthe prepolymer composition, the magnitude of pressure developed in theimplants as it polymerizes, as well as other aspects. If the implant hasrelaxed to an undesirable degree, the clinician has the option to injectadditional implant over the existing implant, thereby achieving a lesscompliant implant. But in general, the implant will establish atherapeutically beneficial disc height that balances deformation of theendplates (axial force) with deformation of the annulus (radial force).The initial ratio of radial to axial forces and the time progression ofthis ratio are selected by the medical professional according to theparadigm outlined in the Appendix.

Once this equilibrium is achieved in the patient, which can additionallydepend on the confluence of acute and chronic forces and the way inwhich the implant responds to these forces, there is a need for theimplant to become less accommodative. This can be achieved naturally bythe implant as the bubbles in the implant fill in with water, therebyincreasing the Poisson ratio of the implant. This infiltration ofbubbles with water fixes both the equilibrated disc height and ratio ofaxial and radial forces. The principal change precipitated by thefilling of the implant bubbles is the implant's accommodative nature.

In one embodiment, liquid implants curing to a solid hydrogel with agaseous fraction possess some degree of level accommodation as the gasin the bubbles is dissolved in the aqueous phase and water migrates tofill the bubbles of the implant and a mean disc height is established.These implants will possess a Poisson ratio less than 0.5. The Poissonratio can be controlled independently of the fraction of gas liberatedby controlling the pressure at which the implant cures. In oneembodiment, prepolymers with a % NCO of between 2 and 5 provide adequategas production to be Poisson ratio adjustable to suit the treatmentrange by applying pressure. In another embodiment, low-pressureapplications, the % NCO ranges from 1.0 to 2.5%.

Another embodiment relates to the ability of the implant to distributeload forces between axial and radial components. This distribution offorces occurs acutely as the implant polymerizes and chronically as itresponds to the unique conditions of the implant environment.Accordingly, the implants tend to have a lower modulus than the nuclearimplants of the past. This embodiment makes maximal use of anystructural strength existing in the affected disc both duringimplantation and subsequently. For example, to minimize the force perunit area applied to the endplates the implant fills the nuclear spaceand distributes localized forces. Similarly the annulus can be made tobear more of the load when the implant applies radial forces to theannulus, making the annulus stiffer in the axial direction. In most discdisease states, the normal equal allocation of load to the centralportions of the endplates and the annulus is compromised, with most ofthe load directed to the center of the endplates. Placing a localizedimplant in the nucleus that is deformable and space filling reallocatesto some degree the distribution of load between the central portion ofthe endplates and the annulus. This reallocation can be determined by anumber of factors, some of which depend on the hoop strength of theannulus, equilibrated disc height, and the degree to which the bubblesin the implant are pressurized. While the dynamics of this allocativeprocess are complex, the mechanism for achieving reallocation of loadforces can depends upon the ability of the implant to apply radialforces to the disc annulus.

For suitably soft cured implants, as those detailed in Table 1, axialforces applied to a disc can cause the implant to expand radial, applyforce to the annulus and cause it to stiffen. The increased stiffness ofthe annulus can result in the annulus supporting a greater proportion ofthe total load upon further disc compression. A component of thelevel-seeking feature is that the distribution of forces between thenucleus and annulus is modulated by the application of radial forces bythe nuclear material. An implant of very high modulus will apply noforces to the annulus, and substantially all the load will be supportedby the implant. In one embodiment for a liquid implant, the axial forceand radial force is equal within the implant. For Poisson ratio 0.5implants of moderate modulus, the modulus determines the partition ofaxial and radial forces. For an implant with a Poisson ratio of lessthan 0.5, the volume displacement of the implant in the axial directiondoes not have to equal the volume displacement in the radial direction,and allows for the height at which the implant modulus is applied tovary. This can allow the implant to respond to a broad spectrum ofloading frequencies. The bubbles in the implant absorb higherfrequencies, and accordingly the implant need not dilate as far radialpotentially protecting the implant from fracture. Lower frequency andmean loads set a mean compression of the implant bubbles, the gas inthese bubbles eventually dissolve into the implant as liquid fills thebubbles. When the bubbles are filled, the height of the implant is nowfixed at a mean height and the modulus of the hydrogel component of theimplant dominates because the compressibility of the implant goes tozero. At this established height, the load and the modulus of theimplant determine the axial and radial dimensions of the implant.

One embodiment provides a liquid nucleus implant sufficient to providethe therapeutic effect of strengthening and/or filling theintervertebral space and preventing extrusion of the polymerizedprosthetic. Although a variety of in-situ polymerizing liquids may beused, both adhesive and non-adhesive, an exemplary in-situ polymerizingliquid is a single-component polyisocyanate based adhesive as describedin U.S. Pat. Nos. 6,254,327, 6,296,607, and U.S. Pub. Nos. 2005/0129733and 2005/0215748, the disclosures of which are incorporated herein byreference.

In one embodiment, the prepolymer comprises a polyisocyanate-cappedpolymeric polyol and a small amount of free poly isocyanate (e.g., 1% to5% by weight). Such materials and their synthesis are described indetail in U.S. Pat. No. 6,524,327, the disclosure of which isincorporated herein by reference. The small amount of excesspolyisocyanate, typically of molecular weight less than about 1000Daltons, maximizes the reactivity of the polyols, and by directly andrapidly reacting with tissue, promotes bonding of the adhesive totissue. Typically the small isocyanate contains up to about 3% of thenumber of active isocyanate groups on the polymer. The small isocyanatemay be all or part low molecular weight capped diol. The capped polyolis multifunctional, and typically is trifunctional or tetrafunctional,or a mixture of trifunctional and/or tetrafunctional with difunctional.The polyol is can be at least in part a polyether polyol.

The polyisocyanate can be difunctional, tri- or tetrafunctional, or starforms of isocyanate. Branching (tri- or tetra-functionality) may beprovided by a trifunctional polymer, or by providing a tri- ortetrafunctional low molecular weight polyol, such as glycerol,erthyritol or isomer, or trimethylolpropane (TMP). Fast reactingformulations use an aromatic diisocyanate such as toluene diisocyanate.Slow reacting formulations use an aliphatic diisocyanate such asisophorone diisocyanate. Many additional diisocyanates are potentiallyuseful. Some are listed in U.S. Pat. No. 6,524,327, and these and othersare found in chemical catalogs, for example from Aldrich Chemical.Alternatively, the polymerization time can be adjusted by selectingappropriate molecular weight polyols. The higher molecular weightpolyols produce lower viscosity capped reaction products and slowerreacting solutions. However, at any molecular weight of the polyol(s),the reaction rate is most significantly determined by the reactivity ofthe functional end group attached to the polyol.

In one embodiment, the prepolymer is an adhesive prepolymer. In otherembodiments, nonadhesive prepolymers can be used. An adhesive for use inthe invention can be hydrophilic in character, and can also bewater-soluble before being crosslinked. This hydrophilicity enable theadhesive to be injected into tissue to polymerize in contact with, andbond to, the tissue, as adhesive and/or as local bulking agent to fillgaps or fissures, or to stabilize implants. The adhesive acts as aself-sealing fluid when injected into cavities or gaps. Once cured insitu, the hydrophilic adhesive will absorb fluid from the tissue,forming a structure that will be at least somewhat gel-like incharacter. The cured adhesive will swell to a controlled extent,exerting a controlled amount of local pressure.

The tensile properties of the cured adhesive can be adjusted so that theadhesive, like the native tissues of the annulus or of the nucleus,deforms under pressure while exerting a restorative force on thesurrounding structures. Hence, the adhesive-tissue composite tends toreturn to its original shape and location after movement of the spineand is characteristically elastic. These properties can be controlled bythe composition of the adhesive, or by providing a controlled degree ofdilution with saline at the time of administration. This is in contrastwith rigid materials, which tend to fracture rather than yield, and toflowable media, which have no tendency to return to their original shapeafter relaxation of stress. For example, hydrophobic adhesives tend tobecome rigid, favoring fracture of the cured adhesive at the surface ofthe tissue or implant. They also tend not to bond to tissue, which ishighly hydrophilic.

Polymeric compositions other than isocyanate-capped polyols can besuitable. The cured implant may be stable in the body, or may degrade inthe body to smaller, excretable molecules (“degradable”). A wide varietyof linkages are known to be unstable in the body. These include, withoutlimitation, esters of hydroxy acids, e.g., alpha and beta hydroxycarboxylic acids; esters of alpha and beta amino acids; carboxylic acidanhydrides; phosphorous esters; and certain types of urethane linkages.Generally, the cured implant is stable in the body for prolongedperiods, as the fibrous materials of the annulus have very limitedself-repair capabilities, and the nucleus has virtually none. However,if methods are found to enhance natural biological repair of the nucleusor annulus, then degradable adhesives or fillers could be used.

The prepolymers can have reactive groups covalently attached to them, orpart of the backbone. The reactive groups are suitable for reaction withtissue, and for crosslinking in the presence of water or components ofbodily fluids, for example water and protein. Suitable groups includeisocyanate, isothiocyanate, anhydrides and cyclic imines (e.g.,N-hydroxy succinimide, maleimide, maleic anhydride), sulfhydryl,phenolic, polyphenolic, and polyhydroxyl aromatic, and acrylic or loweralkyl acrylic acids or esters. Such reactive groups are most commonlybonded to a preformed polymer through suitable linking groups in thepolymer. Commonly found linking groups include, without limitation,amines, hydroxyls, sulfhydryls, double bonds, carboxyls, aldehydes, andketone groups. Of these groups, aliphatic hydroxyls are among the mostwidely used.

Thus, suitable base polymers include poly(alkyl)acrylic acids andpolyhydroxyalkyl acrylates, polysaccharides, proteins, polyols,including polyetherpolyols, polyvinyl alcohol, and polyvinylpyrrolidone,and these same structures with amine or sulfur equivalents, such aspolyethyleneimine, aminosugar polymers, polyalkylamine substitutedpolyethers, and others. Any of these polymers can be substituted withtwo or three reactive groups, as is required to form a crosslinkablepolymer. When there are many substitutable linking groups, as withpolysaccharides, only a few of the substitutable groups (here, mostlyhydroxyls) should be substituted, and the derivatized polymer will havea somewhat random substitution. In one embodiment, the hydrophilicpolymer will have only a few substitutable linking groups. Polyetherpolyols grown on glycol or amine starters will typically have reactivegroups only at the end of the polyether chains, allowing for detailedcontrol of stoichiometry. In one embodiment, the base polymer is apolymer of ethylene glycol, or a copolymer of ethylene glycol with oneor more of propylene glycol, butylene glycol, trimethylene glycol,tetramethylene glycol, and isomers thereof, wherein the ratio ofethylene glycol to the higher alkanediol in the polymer is sufficient toprovide substantial water solubility at room or body temperature. Suchpolymer substrates can be synthesized by known methods. Preformedpolyetherpolyols can be purchased, optionally in a prequalified medicalgrade, from any of numerous catalogs or manufacturers.

The prepolymer can be liquid at room temperature (ca. 20° C.) and bodytemperature (ca. 37° C.), for ease of administration and of mixture withadditives, etc. The prepolymer can be stable in storage at roomtemperature, when protected from moisture and light.

The prepolymer may be supplemented by the addition, during manufactureor at the time of administration, of ancillary materials. These mayinclude reinforcing materials, drugs, volume or osmotic pressurecontrolling materials, and visualization aids for optical, fluoroscopicultrasound or other visualization of fill locations. Reinforcingmaterials may include particulate materials, fibers, flocks, meshes, andother conventionally used reinforcers. These be commercial materials canbe approved for in vivo medical use. Visualization materials include awide variety of materials known in the art, such as, among others, smallparticles of metals or their oxides, salts or compounds for fluoroscopy,gas-filled particles for ultrasound, and dyes or reflecting particlesfor optical techniques.

Osmotic properties can be adjusted for immediate or long-term effects.For example, polyether polyol isocyanates have little ionic chargeeither before or after polymerization. However, in some situations, asdescribed below, it may be desirable to have a controlled degree ofswelling in water after curing. This can be controlled in part by theratio of ethylene glycol to other polyols in the formulation. It canalso be adjusted by adding charged groups to the formulation. A simplemethod is to add charged polymers or charged small molecules to theadhesive at the time of application, for example dissolved in an aqueoussolution. Charged polymers, such as polyacrylic acids, will react poorlywith the isocyanates, but will tend to be trapped in the polymerizedmatrix. They will tend to increase the swelling of the cured material.In turn, this would allow the use of higher proportions of non-ethyleneglycol monomers in the polyols. Alternatively, charge could beintroduced by addition of hydroxy carboxylic acids, such as lactic acid,or tartaric acid, during synthesis or during administration. Addedpolymers could instead be polyamines, but, to avoid rapidpolymerization, should be tertiary or quaternary amines or other aminetypes that will not react with isocyanate. A method of increasingswelling is to incorporate higher concentration of diffusible ions, suchas soluble salts—e.g., sodium chloride—into the adhesive at the time ofapplication. The salt will attract water into the adhesive polymers;after polymerization, the salt will diffuse away and the gel will remainexpanded.

The prepolymer can be adjusted in several ways to optimize its post-cureproperties for the particular situation. In one embodiment, adjustmentof properties is achieved by dilution of the polymer with water, saline,or other aqueous solution. A typical dilution would be in the range of5% or less (volume of saline in liquid polymer), for formation of dense,high-tensile, low-swelling deposits, up to about 95% (19 vol.saline/vol. polymer) for readily swelling, highly compliant deposits orbonds. In formulation, allowance must be made for the amount of waterthat will flow into the adhesive from the tissue during reaction. Thiswill usually be relatively small for bulk deposits, but is of moreconcern for thin adhesive layers. In thin layers, fast-curingcompositions can be used, such as compositions with a higher proportionof aromatic diisocyanates. In general, dilution will reduce the tensilestrength and the modulus. The amount of dilution will tend to bedifferent depending on whether the modulus or tensile strength is tomatch that of the annulus (higher) or the nucleus (lower).

Various non-reactive ingredients can be added to the polymer solutioneither in the prepolymer or in the aqueous solution to alter thehydrogel mechanical properties, e.g., tensile strength, elasticity andbubble size. Inert particulate such as tantalum powder will result inbubble nucleation and a finer bubble size, increase the modulus of thehydrogel, and make the hydrogel radio opaque. Emulsifiers can be addedto increase mix homogeneity, reduce bubble size, and provide a higherelongation at break. It is possible to use the same diol used toconstruct the prepolymer as an emulsifier. Alternatively, a higher orlower molecular weight diol may be used. The ratio of EO/PO can bealtered to increased mixability, or pure forms of EO or PO can be used.When pure EO is used, the mixture of prepolymer and aqueous solutionbecomes non-Newtonian, and tends to take on a stringy consistency, whichcan further improve elasticity.

Other adjustable factors include the molecular weight of the polymer,and its degree of branching; and its hydrophilicity, which is a functionof the particular polyol or polyols used in the formulation. Inaddition, additives, as described above, can also influence theseproperties.

Polymeric Compositions

Disclosed herein are liquid preparations for use in medicine. The liquidpreparation contains a reactive polymer, which comprises a “basepolymer” or “backbone polymer”, reactive groups on the backbone polymer,and a slight excess of “free” (low molecular weight) polyreactivemolecules. The liquid composition can be prepared by a method requiringno catalysts and essentially no solvent. The reactive liquid polymer isself-curing when applied to tissue, by absorption of water and otherreactive molecules from the tissue. The cured polymer can seal tissue totissue, or to devices; to apply a protective coating to tissue; to forman implant within or upon tissue; to deliver drugs. The cured polymercan be provided with biodegradable groups, and has a controllable degreeof swelling in bodily fluids.

Backbone Polymers

In one embodiment, the backbone polymer comprises a polymeric segment,of molecular weight about 500 D or more, e.g., about 1000 to about10,000 D, optionally up to about 15 kD or 20 kD. The backbone polymercan contain groups that can be easily derivatized (“capped”) to form thefinal reactive group. Such groups can be alcohols or amines, oroptionally sulfhydryls or phenolic groups. Examples include polymerssuch as a polymeric polyol, or optionally a polymeric polyamine orpolyamine/polyol. In one embodiment, the polyols are polyether polyols,such as polyalkylene oxides (PAOs), which may be formed of one or morespecies of alkylene oxide. The PAO, when comprising more than onespecies of alkylene oxide, may be a random, block or graft polymer, or apolymer combining these modes, or a mixture of PAO polymers withdifferent properties. Exemplary alkylene oxides are ethylene oxide andpropylene oxide. Other oxiranes may also be used, including butyleneoxide. PAOs are typically made by polymerization onto a startermolecule, such as a low molecular weight alcohol or amine, e.g., apolyol. Starting molecules with two, three, four or more derivatizablealcohols or other derivatizable groups can be used. The multi-armed PAOsobtained from such starters will typically have one arm for each groupon the starter. PAOs with two, three or four terminal groups can beused. Mixtures of PAOs or other backbone polymers, having variablenumbers of arms and/or variation in other properties can be used.

Common polyols useful as starters are aliphatic or substituted aliphaticmolecules containing a minimum of 2 hydroxyl or other groups permolecule. Since a liquid end product is desired, the starters can be oflow molecular weight containing less than 8 hydroxyl or other groups.Suitable alcohols include, for illustration and without limitation,adonitol, arabitol, butanediol, 1,2,3-butanetriol, dipentaerythritol,dulcitol, erythritol, ethylene glycol, propylene glycol, diethyleneglycol, glycerol, hexanediol, iditol, mannitol, pentaerythritol,sorbitol, sucrose, triethanolamine, trimethylolethane,trimethylolpropane. Small molecules of similar structures containingamines, sulfhydryls and phenols, or other groups readily reactive withisocyanates, are also useable.

The PAO, or other backbone polymer, may optionally incorporate non-PAOgroups in a random, block or graft manner. Non-PAO groups are optionallyused to provide biodegradability and/or absorbability to the finalpolymer. Groups providing biodegradability are well known. They includehydroxy carboxylic acids, aliphatic carbonates, 1,4-dioxane-2-one(p-dioxanone), and anhydrides. The hydroxy carboxylic acids may bepresent as the acid or as a lactone or cyclic dimmer, and include, amongothers, lactide and lactic acid, glycolide and glycolic acid,epsilon-caprolactone, gamma-butyrolactone, and delta-valerolactone.Amino acids, nucleic acids, carbohydrates and oligomers thereof can beused to provide biodegradability. Methods for making polymers containingthese groups are well known, and include, among others reaction oflactone forms directly with hydroxyl groups (or amine groups),condensation reactions such as esterification driven by water removal,and reaction of activated forms, such as acyl halides. Theesterification process involves heating the acid under reflux with thepolyol until the acid and hydroxyl groups form the desired ester links.The higher molecular weight acids are lower in reactivity and mayrequire a catalyst making them less desirable.

The backbone polymers may also or in addition carry amino groups, whichcan likewise be functionalized by polyisocyanates. Thus, the diaminederivative of a polyethylene glycol could be used. Low molecular weightsegments of amine containing monomers could be used, such asoligolysine, oligoethylene amine, or oligochitosan. Low molecular weightlinking agents, as described below, could have hydroxyl functionality,amine functionality, or both. Use of amines will impart charge to thepolymerized matrix, because the reaction product of an amine with anisocyanate is generally a secondary or tertiary amine, which may bepositively charged in physiological solutions. Likewise, carboxyl,sulfate, and phosphate groups, which are generally not reactive withisocyanates, could introduce negative charge if desired. A considerationin selecting base polymers, other than PAOs or others that react only atthe ends, is that the process of adding the reactive groups necessarilyrequires adding reactive groups to every alcohol, amine, sulfhydryl,phenol, etc. found on the base polymer. This can substantially changethe properties, e.g., the solubility properties, of the polymer afteractivation.

Reactive Groups

The base or backbone polymer is then activated by capping with lowmolecular weight (LMW) reactive groups. In one embodiment, the polymeris capped with one or more LMW polyisocyanates (LMW-PIC), which aresmall molecules, typically with molecular weight below about 1000 D,more typically below about 500 D, containing two or more reactiveisocyanate groups attached to each hydroxyl, amine, etc of the basemolecule. After reaction of the LMW-PIC with the backbone, each capablegroup of the backbone polymer has been reacted with one of theisocyanate groups of the LMW-PIC, leaving one or more reactiveisocyanates bonded to the backbone polymer via the PIC. The LMW-PIC arethemselves formed by conjugation of their alcohols, amines, etc. withsuitable precursors to form the isocyanate groups. Starting moleculesmay include any of those mentioned above as starting molecules forforming PAOs, and may also include derivatives of aromatic groups, suchas toluene, benzene, naphthalene, etc. The LMW-PIC for activating thepolymer can be di-isocyanates, e.g., toluene diisocyanate (TDI) andisophorone diisocyanate, both commercially available. When adiisocyanate is reacted with a capable group on the base polymer, one ofthe added isocyanates is used to bind the diisocyanate molecule to thepolymer, leaving the other isocyanate group bound to the polymer andready to react. As long as the backbone polymers have on average morethan two capable groups (hydroxyl, amine, etc.), the resultingcomposition will be crosslinkable.

A wide variety of isocyanates are potentially usable as LMW-PICs.Suitable isocyanates include 9,10-anthracene diisocyanate,1,4-anthracenediisocyanate, benzidine diisocyanate, 4,4′-biphenylenediisocyanate, 4-bromo-1,3-phenylene diisocyanate, 4-chloro-1,3-phenylenediisocyanate, cumene-2,4-diisocyanate, cyclohexylene-1,2-diisocyanate,cyclohexylene-1,4-diisocyanate, 1,4-cyclohexylene diisocyanate,1,10-decamethylene diisocyanate, 3,3′dichloro-4,4′biphenylenediisocyanate, 4,4′diisocyanatodibenzyl, 2,4-diisocyanatostilbene,2,6-diisocyanatobenzfuran, 2,4-dimethyl-1,3-phenylene diisocyanate,5,6-dimethyl-1,3-phenylene diisocyanate, 4,6-dimethyl-1,3-phenylenediisocyanate, 3,3′-dimethyl-4,4′diisocyanatodiphenylmethane,2,6-dimethyl-4,4′-diisocyanatodiphenyl,3,3′-dimethoxy-4,4′-diisocyanatodiphenyl, 2,4-diisocyantodiphenylether,4,4′-diisocyantodiphenylether, 3,3′-diphenyl-4,4′-biphenylenediisocyanate, 4,4′-diphenylmethane diisocyanate, 4-ethoxy-1,3-phenylenediisocyanate, ethylene diisocyanate, ethylidene diisocyanate,2,5-fluorenediisocyanate, 1,6-hexamethylene diisocyanate, isophoronediisocyanate, lysine diisocyanate, 4-methoxy-1,3-phenylene diisocyanate,methylene dicyclohexyl diisocyanate, m-phenylene diisocyanate,1,5-naphthalene diisocyanate, 1,8-naphthalene diisocyanate, polymeric4,4′-diphenylmethane diisocyanate, p-phenylene diisocyanate,4,4′,4″-triphenylmethane triisocyanate, propylene-1,2-diisocyanate;p-tetramethyl xylene diisocyanate, 1,4-tetramethylene diisocyanate,toluene diisocyanate, 2,4,6-toluene triisocyanate, trifunctional trimer(isocyanurate) of isophorone diisocyanate, trifunctional biuret ofhexamethylene diisocyanate, and trifunctional trimer (isocyanurate) ofhexamethylene diisocyanate.

In general, aliphatic isocyanates will have longer cure times thanaromatic isocyanates, and selection among the various availablematerials will be guided in part by the desired curing time in vivo. Inaddition, commercial availability in grades suitable for medical usewill also be considered, as will cost. Toluene diisocyanate (TDI) andisophorone diisocyanate (IPDI) can be used. The reactive chemicalfunctionality of the liquids can be isocyanate, but may alternatively orin addition be isothiocyanate, to which all of the above considerationswill apply.

Physical Properties of the Cured Implant

The polymerizable materials are typically liquids at or near bodytemperature (i.e., below about 45° C.), and can be liquid at roomtemperature, ca. 20-25° C., or below. The liquids are optionallycarriers of solids. The solids may be biodegradable or absorbable. Theliquid polymerizable materials are characterized by polymerizing uponcontact with tissue, without requiring addition of other materials, andwithout requiring pretreatment of the tissue, other than removing anyliquid present on the surface(s) to be treated. A related property ofthe polymerizable materials is that they are stable for at least 1 yearwhen stored at room temperature (ca. 20-25 degrees C.) in the absence ofwater vapor. This is because the material has been designed so that boththe reaction that polymerizes the polymers, and the reactions thatoptionally allow the polymer to degrade, both require water to proceed.

In contrast to previous formulations, the polymeric polyisocyanatescontain a low residual level of low molecular weight (LMW)polyisocyanates (PIC). For example, the final concentration of LMW-PICisocyanate groups in the formulation, expressed as the equivalentmolarity of isocyanate groups attached to LMW compounds, is normallyless than about 1 mM (i.e., 1 mEq), e.g., less than about 0.5 mEq, oreven less than about 0.4 mEq. In one embodiment, the level of LMWisocyanate groups is finite and detectable, for example greater thanabout 0.05 mEq, or greater than about 0.1 mEq. It is believed thathaving a low but finite level of LMW-PIC molecules tends to promoteadherence between the applied polymer formulation and the tissue beingtreated. However, decreased levels of LMW-PIC may tend to decreasetissue irritation during application and cure of the liquid polymerpreparation. It is believed that the range of about 1 mEq to about 0.05mEq is approximately optimal. In situations requiring tissue adherencein the presence of significant biological fluid, or in adherence todifficult tissues, greater levels of LMW-PIC isocyanate groups may beused.

Exemplary Polymer Structures

There are several ways in which the above-recited steps can be used toobtain a suitable liquid reactive polymer system. In a simple system, apolymeric polyol with a number of end groups on average greater than twois treated with a slight excess of a LMW-PIC, such as toluenediisocyanate. The reaction product is formed under nitrogen with mildheating, e.g., by the addition of the LMW-PIC to the polymer. Theproduct is then packaged under nitrogen, typically with no intermediatepurification.

An exemplary biodegradable polyol composition includes a trifunctionalhydroxy acid ester (e.g., several lactide groups successively esterifiedonto a trifunctional starting material, such as trimethylolpropane, orglycerol). This is then mixed with a linear activated polyoxyethyleneglycol system, in which the PEG is first capped with a slight excess ofa LMW-PIC, such as toluene diisocyanate. Then mixing together theactivated polyoxyethylene glycol and the lactate-triol forms theactivated polymer. Each lactate triol binds three of the activated PEGmolecules, yielding a prepolymer with three active isocyanates at theend of the PEG segments, and with the PEG segments bonded togetherthrough degradable lactate groups. In the formed implant, the lactateester bonds gradually degrade in the presence of water, leavingessentially linear PEG chains that are free to dissolve or degrade.Interestingly, in this system, increasing the percentage of degradablecrosslinker increases rigidity, swell and solvation resistance in theformed polymer.

Other polyol systems include hydroxy acid esterified linear polyetherand polyester polyols optionally blended with a low molecular weightdiol. Similarly, polyester and polyether triols esterified with hydroxyacid are useful. Other polyol systems include the use of triol formingcomponents such as trimethylolpropane to form polyols having three armsof linear polyether chains.

Exemplary materials are described in U.S. Pat. No. 6,254,327, and U.S.Pub. Nos. 2003/0135238 and 2004/0068078, the disclosures of which areincorporated herein by reference.

These embodiments provides methods and compositions for treatingintervertebral disc disorders by providing a disc nucleus implant havingone or more of the following: 1) the cured implant comprises solid,liquid and gas phases, 2) the cured implant fills the entire volume ofan implantation formed in a vertebral disc, 3) the liquid implant bondsto tissue as it cures, 4) the liquid implant is of sufficiently lowviscosity to be delivered by conventional means via a tube directed tothe implantation site, 5) the liquid implant provides a cure time shortenough to ensure the liquid implant is localized to an intendedimplantation site and long enough to allow the implant to be deliveredto the site clinically, 6) the cured implant does not change volumebeyond a therapeutic range, 7) the liquid implant releases a gas phasewhile curing that acts to pressurize the implantation site as theimplant cures, 8) the liquid implant contains a radio-opaque orilluminating marker, 9) the cured implant possesses a bimodularcompliance, 10) the compliance of the cured implant can be adjusted in apredictable manner by the medical professional, 11) the liquid implantseals against tissue surfaces within the disc and at the implantationaccess to prevent loss of the native nucleus pulposus or implant whileit cures, 12) the cured implant can exchange its water phase withsurrounding tissue, 13) the cured implant allows nutrient diffusionthrough the implant, 14) the cured implant is clinically permanent or atleast resides functionally useful in the body longer than otherdescribed nucleus implants, 15) the liquid implant is comprised of onefunctional part, 16) the functional part of the liquid implant compriseschemical species that are stable when store together and the minimummolecular weight is large enough to improve biocompatibility, 17) theliquid implant is hydrophilic, 18) the liquid implant bonds to existingimplant, 19) the cured implant is clinically removable in one piece, 20)the liquid implant provides clinically significant reinforcement of thedisc annulus when cured, 21) the liquid implant does not require amolding element or balloon in order to be safe and effective, 22) theliquid and cured implant is not neurotoxic, and more generally isbiocompatible, 23) the cured implant possesses a biocompatibility andmodulus that reduces the chance of implant degradation and migrationthrough the annulotomy, 24) the liquid implant is self-adjusting withinthe body so as to result in a therapeutic volume that makes the clinicaloutcome insensitive to method of implantation, and 25) the cured implanttranslates axial forces originating at the vertebral endplates intoradial forces applied to the disc annulus.

Another embodiment provides the use of preformed foams, which isdistinct from the use of an in situ curable implant material. Thepreformed foams can have one or more of the same properties as the curedfoam materials as described herein. These foams are sufficientlycompressible to allow implantation through an opening having dimensionssmaller than that of the disc space. Exemplary preformed foams includethe nonabsorbable foams comprising polyurethanes,polytetrafluoroethylene, silicone foams, epoxies, and polyvinyl chloridefoam.

One embodiment provides a method of repairing a defect in a spinal discspace, comprising:

inserting a nonabsorbable, closed cell foam having a Poisson ratio ofless than 0.5 into the defect.

The inserting can be performed with a delivery catheter having a lumenof sufficient large diameter for inserting a preformed foam. In oneembodiment, prior to the inserting, the method further comprisesremoving some or all of the nucleus pulposus within the spinal discspace, and the inserting results in replacement of the removed nucleuspulposus with the foam.

In one embodiment, the removing further comprises removing portions ofthe annulus fibrosus in the vicinity of the nucleus pulposus.

EXAMPLES Example 1 Preparation of Prepolymer

In this example an isocyanate terminated diol is trifunctionalized toyield a slow curing tissue adhesive. The type and amount of isocyanateto be used is 326.27 g of isophorone diisocyanate (IPDI). A suitableIPDI is Desmodur I. The type and amount of diol to be used is 749.94 gof 75:25 diol comprised of 75% polyethylene glycol and 25% polypropyleneglycol. A suitable diol is Ucon 75-H-450, with a molecular weight of 978Daltons and hydroxyl number of 119.4. The type and amount of triol to beused is 23.79 g of trimethylol propane. The theoretical target forcompletion of the diol termination steps is % NCO=5.23%. The theoreticaltarget for completion of the trifunctionalization step is % NCO=3.09%.Final temperature pre-TMP was 80° C. The NCO levels at various timesare: at 28 hrs 6.197%, at 56 hrs 5.468%, at 78 hrs 5.421, and at 126 hrs5.23%. The TMP was added at hour 127. The final NCO of % NCO=3.09% wasreached at hour 271. The viscosity at 34° C. was 103 Kcps.

The TMP and glycols should be deionized and dried. All of the diol andisocyanate are to be added at once.

The temperature in the reacting chamber should follow the schedulesdescribed above, and the % NCO at the described time points shouldfollow the values recorded above. The reaction should be conducted undervacuum with a trickle flow of argon.

The reactor is a standard cylindrical glass 1 Liter reactor with a stirrod comprising 2 reactor blades of 55 mm diameter with 5 blades oriented45° from the axis. The rate of mixing is 220 rpm.

Under these conditions the prepolymer is comprised of a broaddistribution of chain lengths in the diol termination phase with aminimum of side reactions. This distribution cannot be achieved solelyby adding diols of molecular weights in the ratio obtained in the finalsynthesis product, since the actual synthesis process can affect thefinal chain length distribution. Adding the diols in this ratio at thebeginning of the synthesis process results in a prepolymer that isunusable as a tissue adhesive. Calling the single chain length of 978Dalton the monomer, the following distribution is obtained after thediol termination process.

Actual Value Useful Range Monomer 28.2% +/−10%  Dimer 20.0% +/−10% Trimer 14.8% +/−5% Tetramer 10.9% +/−2% Pentamer 9.9% +/−1% Hexamer 6.8%+/−1% Heptamer 5.2% +/−1% Octamer 3.3% +/−1% Nonamer 1.8% +/−0.5%  

Example 2 Prepolymer Synthesis

The synthesis is the same as Example 1, except that the diol is added in1% increments rather than all at once to the isocyanate. Each 1%increment of diol added to the reacting isocyanate is made after theexotherm of the previous addition is complete. This step-wise additionyields the following distribution of terminated diols:

Actual Value Useful Range Monomer 55.3% or more Dimer 27.1% or lessTrimer 8.5% or less Tetramer 4.7% or less Pentamer 2.5% or less Hexamer1.3% or less Heptamer 0.6% or less

Example 3 Prepolymer

In this example an isocyanate terminated diol is trifunctionalized toyield a fast curing tissue adhesive. Fast adhesives cure within 5minutes when used neat and applied to tissue. Slow adhesives cure afterthis time, generally 5 to 10 times longer. The type and amount ofisocyanate to be used is 270.26 g of toluene diisocyanate (TDI). Asuitable TDI is Rubinate, a mixture of 80% 2-4 and 20% 2-6 isomers. Thetype and amount of diol to be used is 870.53 g of Ucon 75-H-450. Thetype and amount of triol to be used is 9.21 g of trimethylol propane.The theoretical target for completion of the diol termination steps is %NCO=4.55%. The theoretical target for completion of thetrifunctionalization step is % NCO=3.76%. Final temperature pre-TMP was50° C. The NCO levels at 25 hrs 4.78% and at 75 hrs 4.55%. Then the TMPwas added at hour 76. The final NCO of % NCO=3.67% was reach at hour100. The viscosity at 31° C. was 24.5 Kcps.

Example 4 Prepolymer

In this example an isocyanate terminated diol is trifunctionalized toyield a fast curing tissue adhesive with a ratio of soft-to-hard centersgreater than that achieved in Example 3. The type and amount ofisocyanate to be used is 231.65 g of toluene diisocyanate (TDI). Thetype and amount of diol to be used is 870.53 g of Ucon 75-H-450. Thetype and amount of triol to be used is 9.21 g of trimethylol propane.The theoretical target for completion of the diol termination steps is %NCO=3.90%. The theoretical target for completion of thetrifunctionalization step is % NCO=2.69%. Final temperature pre-TMP was50° C. The NCO levels at 23 hrs 3.80% and at 75 hrs 4.55%. Then the TMPwas added at hour 23. The final NCO of % NCO=2.69% was reach at hour 72.The viscosity at 30° C. was 48 Kcps.

Example 5 Prepolymer

In this example two isocyanate terminated diols are randomlytrifunctionalized to yield a fast curing, absorbable tissue adhesive.The type and amount of isocyanate to be used is 270.26 g of toluenediisocyanate (TDI). The types and amounts of diol to be used is 870.53 gof Ucon 75-H-450 and 25 g poly(DL-lactide-co-glycolide) (50:50). Theaverage molecular weight of the copolymer is 50,000 Dalton. The type andamount of triol to be used is 9.21 g of trimethylol propane. Thetheoretical target for completion of the diol termination steps is %NCO=4.55%. The theoretical target for completion of thetrifunctionalization step is % NCO=3.00%. Final temperature pre-TMP was50° C. The NCO levels at 96 hrs 4.99% and at 312 hrs 4.41%. Then the TMPwas added at hour 312. The final NCO of % NCO=2.93% was reach at hour528. The viscosity at 32° C. was 240 Kcps.

Example 6 Prepolymer

In this example a high molecular weight diol is terminated and randomlytrifunctionalized to yield a slow curing, low viscosity implant. Thetype and amount of isocyanate to be used is 171.29 g of isophoronediisocyanate (IPDI). The type and amount of diol to be used is 824.93 gof Ucon 75-H-1400. The molecular weight of 75-H-1400 is 2500 Dalton. Thetype and amount of triol to be used is 12.49 g of trimethylol propane.The theoretical target for completion of the diol termination steps is %NCO=3.3%. The theoretical target for completion of thetrifunctionalization step is % NCO=2.2%. Final temperature pre-TMP was80° C. The NCO levels at 168 hrs 4.54% and at 624 hrs 3.32%. Then theTMP was added at hour 625. The final NCO of % NCO=2.2% was reach at hour824. The viscosity at 32° C. was 150 Kcps.

Example 7 Prepolymer

In this example a high molecular weight diol is terminated and randomlytrifunctionalized to yield a fast curing, low viscosity tissue adhesive.The formula for Example 6 is used substituting molar equivalents of TDI.

Example 8

In this example, any of Examples 1-7 where the triol, TMP, issubstituted with a molar equivalent of TONE polyol 0301 manufactured byUnion Carbide. The molecular weight of this triol is 300 Dalton with ahydroxyl number of 560.

Example 9 Liquid Prosthetic

In some medical applications a tissue bonding adhesive that does notappreciably swell during polymerization is useful. Applications includedisc nucleus replacement, disc annulus augmentation, and any applicationwhere large static forces predominate. For these applications anadhesive of low % NCO can be used. It is also advantageous to initiatepolymerization outside the body by pre-mixing the tissue adhesive withwater. The amount of water added determines cure time and cured modulus.A useful adhesive for these applications can be prepared by mixingExample 7 in the following ratios with water.

% prepolymer % water cure time Modulus 33 67 1 minute disc nucleus like1-2 N/cm² 50 50 2 minutes disc annulus like 2-3 N/cm² 70 30 3 minuteshardest 3-4 N/cm²

Example 10 Liquid Prosthetic

The cured modulus of an adhesive can be increased by adding aparticulate. When 0.3 micron tantalum powder is added, the material canbe made radio-opaque. A higher modulus implant can be made by adding 10%by volume tantalum powder to the mixtures of Example 9.

Example 11

Two human lumbar spine specimens were utilized in a study of loadfailure. The study compared nucleus replaced by a gel vs nucleusreplaced by a foam of Example 3. The specimens were screened for grossanatomical defects. The age of the donors did not exceed 85 years.

TABLE 2 Specimen Levels Used L4/5 L2/3 T12/L1 L3/L4 L = Lumbar, T =Thoracic

The specimens were thawed to room temperature and all residualmusculature removed via careful dissection. Throughout preparation andtesting, the specimens were kept moist with a wrapping of saline-soakedgauze. A total of four two-level spinal lumbar spine segments wereharvested. To create Anterior Column Units, the pedicles were transectedand the posterior elements removed. Care was taken to preserve allremaining ligamentous attachments and maintain segmental integrity. Foreach segment, the cephalad and caudad vertebrae were rigidly embedded ina urethane potting compound. The segments were potted so that themid-plane of the intervertebral disc was horizontal. Sufficient spacewas left for injection of the nucleus replacement device (Gel or Foam).

The spinal segments were tested using a standard compression protocoldeveloped by Rhode Island Hospital with custom fixtures in an MTS 810servohydraulic load frame. The upper compression platen was not allowedto rotate. The segments were loaded to failure in compression at a rateof 0.167 mm/sec. The mode of failure and maximum load were recorded.

Implant Type Mean Failure (N) Failure Mode Gel 4100 +/− 2400 endplatefracture Foam 7500 +/− 1000 endplate fracture

From the higher mean failure value, it can be seen that the foamimplants having a Poisson ratio <0.5 allows the disc to withstand higherloads without failure compared to preformed liquid/gel implants.

APPENDIX

Ideal physical characteristics of a disc nucleus replacement prostheticsare presented. Emphasis is placed on in situ polymerizing prosthetics,e.g., hydrogel forming prepolymers. One feature is identification of theoptimal range of prosthetic moduli for various disc dimensions andloading conditions that satisfy the derived endplate limit and loaddeflection requirements. Formulae for matching prosthetic moduli tovarious pathological conditions of the annulus fibrosus are disclosed.

Disclosed generally herein is the formation of implants for the repairof lesions in the human spinal disc, especially in its nucleus, andmaterials and material properties for such implants. Disc repairimplants may be either preformed or formed in the body. Also disclosedis the selection of optimal mechanical characteristics for nucleusreplacement materials. Also disclosed are clinically effective rangesfor mechanical properties, e.g., the modulus, for nucleus replacementdevices placed in the nuclear space of a spinal disc with a defectiveannulus fibrosis.

BACKGROUND A. Treatment of Spinal Disc Abnormalities

Intervertebral disc abnormalities are common in the population and causeconsiderable pain, particularly if they affect adjacent nerves. Discabnormalities result from trauma, wear, metabolic disorders and theaging process and include degenerative discs, localized tears orfissures in the annulus fibrosis, localized disc herniation withcontained or escaped extrusions, and chronic, circumferential bulgingdiscs. Disc fissures occur as a degeneration of fibrous components ofthe annulus fibrosis. Rather minor activities such as sneezing, bendingor simple attrition can tear degenerated annulus fibers and create afissure. The fissures may be further complicated by extrusion of nucleuspulposus material into or beyond the annulus fibrosis. Difficulties canstill present even when there is no visible extrusion, due tobiochemicals within the disc irritating surrounding structures andnerves. Initial treatment includes bed rest, painkillers and musclerelaxants, but these measures rarely correct the underlying cause.Surgical treatments include reduction of pressure on the annulus byremoving some of the interior nucleus pulposus material by percutaneousnuclectomy. Surgical treatments meant to cure the underlying causeinclude spinal fusion with screws, rods and fusion cages. Devices andprocedures involving screws, rods and plates are disclosed in thefollowing U.S. patents, as well as others: Errico U.S. Pat. Nos. 37,665;5,733,286; 5,549,608; 5,554,157; 5,876,402; 5,817,094; 5,690,630;5,669,911; 5,647,873; 5,643,265; 5,607,426; 5,531,746 and 5,520,690;Metz-Stavenhagen U.S. Pat. No. 6,261,287; Puno U.S. Pat. No. 5,474,555;Byrd U.S. Pat. No. 5,446,237; Biedermann U.S. Pat. Nos. 5,672,176 and5,443,467; Cotrel U.S. Pat. Nos. 4,815,453 and 5,005,562; Jackson U.S.Pat. No. 5,591,165; Harms U.S. Pat. Nos. 4,946,458; 5,092,867; 5,207,678and 5,196,013; Mellinger U.S. Pat. No. 5,624,442; Sherman U.S. Pat. Nos.5,885,286; 5,797,911 and 5,879,350; Morrison U.S. Pat. No. 5,891,145;Tatar U.S. Pat. No. 5,910,142; Nicholas U.S. Pat. No. 6,090,111; andYuan U.S. Pat. No. 6,565,565. Fusion cages and related procedures aredisclosed in Bagby U.S. Pat. No. 4,501,269; Michelson U.S. Pat. Nos.5,015,247 and 5,797,909; Ray U.S. Pat. No. 6,042,582 and Kuslich U.S.Pat. Nos. 5,489,308; 6,287,343 and 5,700,291. Proposed disc replacementdevices are disclosed in the following U.S. patents: Middleton U.S. Pat.No. 6,315,797; Marnay U.S. Pat. No. 5,314,477; Stubstad U.S. Pat. No.3,867,728; Keller U.S. Pat. No. 4,997,432; and Buettner-Janz U.S. Pat.No. 4,759,766.

A contained disc herniation is not associated with free nucleusfragments migrating to the spinal canal. However, a contained discherniation can still protrude and irritate surrounding structures, forexample by applying pressure to spinal nerves. Escaped nucleus pulposuscan chemically irritate neural structures. Current treatment methodsinclude reduction of pressure on the annulus by removing some of theinterior nucleus pulposus material by percutaneous nuclectomy. See, forexample, Kambin U.S. Pat. No. 4,573,448. Complications include discspace infection, nerve root injury, hematoma formation, and instabilityof the adjacent vertebrae and collapse of the disc from decrease inheight. It has been proposed to treat weakening due to nucleus pulposusdeficiency by inserting preformed hydrogel implants. See, Ray U.S. Pat.Nos. 4,772,287; 4,904,260 and, 5,562,736 and Bao U.S. Pat. No.5,192,326.

Circumferential bulging of the disc also can result in chronic discweakening. The joint can become mechanically less stable. As the bulgingdisc extends beyond its normal circumference, the disc height iscompromised and nerve roots are compressed. In some cases osteophytesform on the outer surface of the disc and further encroach on the spinalcanal and channels through which nerves pass. The condition is known aslumbar spondylosis. Continued disc degeneration can resulting in onevertebral body segment approaching and possibly contacting an adjacentvertebral body segment.

Treatment for segmental instability include bed rest, pain medication,physical therapy and steroid injection. Spinal fusion is the finaltherapy performed with or without discectomy. Other treatment includesdiscectomy alone or disc decompression with or without fusion.Nuclectomy can be performed by removing some of the nucleus matter toreduce pressure on the annulus. Complications include disc spaceinfection, nerve root injury, hematoma formation, and instability ofadjacent vertebrae. New fixation devices include pedicle screws andinterbody fusion cages. Studies on fixation show success rates between50% and 67% for pain improvement, and a significant number of patientshave more pain postoperatively.

Delivery of tissue adhesives to the spine in a minimally invasive mannerhas been disclosed, including procedures for restoring structuralintegrity to vertebral bodies. See Scribner U.S. Pat. Nos. 6,241,734 and6,280,456; Reiley U.S. Pat. Nos. 6,248,110 and 6,235,043; Boucher U.S.Pat. No. 6,607,554 and Bhatnagar U.S. Pat. No. 6,395,007. Methods ofrepairing the spinal disc or portions thereof are disclosed in CauthernU.S. Pat. No. 6,592,625, Haldimann U.S. Pat. No. 6,428,576, Trieu U.S.Pat. No. 6,620,196 and Milner U.S. Pat. No. 6,187,048.

B. Surgical Approaches to the Spine

The spine may be approached in open surgery using posterior, anterior orlateral approaches. The following is a brief description of severalproposed surgical approaches, which may be used to gain access to thespine in a less invasive manner to treat spinal insufficiency.

Posterior Lateral Approach

Methods for disc access include laminectomy, a procedure wherein achannel is made from the dorsal side of the patient's back to the lumbarlamina of the disc. Blood vessels, ligaments, major back support musclesand spinal nerves located around the dural sac are retracted. Once thechannel has been cleared, the standard procedure is to cut a hole in thedisc capsule and pass instruments into the disc interior. This approachcreates a defect that is oriented toward spinal nerves, thus typicallythe nucleus is completely removed to prevent extrusion of nuclearmaterial and subsequent pressure on these nerves. Alternatively, undervisual magnification with an operating microscope or operating loupe,small diameter microsurgical instruments can access the disc withoutcutting bone. It is possible to bypass the nerves and blood vesselsentirely by inserting a cannula through the patient's side above thepelvic crest to reach a predetermined position along the lumbar portionof the spine. This procedure can be guided with use of fluoroscopy.

Kambin U.S. Pat. No. 4,573,448 describes a posterior lateral approachperformed under local anesthesia by the insertion of a cannulated trocarover a guide wire extending through the patient's back toward a targetdisc at an angle of approximately 35 degrees with respect to thepatient's perpendicular line. A hollow needle with a stylet can beinserted at a location spaced from the midline so as to form a 35-degreeangle in an oblique direction. When the needle reaches the annulusfibrosis it is withdrawn after a guide wire is introduced through theneedle to the disc. A cannulated, blunt-tipped trocar is passed over theguide wire until the tip reaches the annulus. The guide wire iswithdrawn. A closely fitting, thin-walled cannula is passed over thetrocar until it reaches the annulus. The trocar can be withdrawn.Cutting instruments or a punch can be used to expose the nucleus.

Paramedian Transabdominal Procedure

In this procedure the patient is in the supine or lithotomy position.This transabdominal procedure involves splitting the paramedian rectus,retracting the bowel, incising the peritoneum on the posterior wall ofthe abdominal cavity and accessing the anterior spine. Alternatively,the anterior rectus sheath is exposed of the left rectus muscle. Theanterior rectus sheath is incised to expose the body of the rectusmuscle. The rectus muscle is then mobilized over an adequate length,preferably symmetrical with the incision, and the rectus is retractedmedially. The posterior rectus sheath is cut to expose the peritoneum.The peritoneum is pushed aside and dissected to expose the psoas muscle.The ureter and the left iliac vessels are mobilized. The rectus muscle,ureter, iliac vessels, and peritoneum are retracted laterally to exposethe lumbar region. For repair to lumbar vertebrae L3-4 and L4-5, accessshould be made to the left of the aorta and inferior vena cava, betweenthe aorta and the psoas muscle, and through the posterior peritoneum andfatty tissue. In some cases it may be necessary to transverse the psoasmuscle. For access to sites between L5 and S-1, the dissection is closerto the midline between the iliac branches of the great vessels.

Lateral Retroperitoneal Procedure

The retroperitoneal procedure involves placing the patient in the rightlateral recumbent position and making an incision in the abdomen at theborder of the rectus muscle and subsequent dissection down to identifythe peritoneum. Dissection can be performed bluntly or may befacilitated using a balloon cannula or expanding cannula as described byBonutti (U.S. Pat. No. 5,514,153). The resulting retroperitoneal cavitycan be held open with a retractor positioned to elevate the wall of thecavity adjacent to the patient's left side. The retractor may be aballoon retractor, see for example Moll U.S. Pat. No. 5,309,896 andBonutti U.S. Pat. Nos. 5,331,975; 5,163,949; 6,277,136; 6,171,236; and5,888,196. The peritoneum is dissected away from the abdominal wall infirst a lateral and then a posterior direction until the spine isexposed. Under endoscopic visualization the iliopsoas muscle isdissected or retracted to facilitate disc repair.

Alternatively, dissection of the peritoneum can be accomplished usinggas pressure into the preperitoneal and retroperitoneal space, therebyexpanding the space and dissecting the peritoneal lining from theabdominal wall while relocating the peritoneal lining toward the midlineof the abdomen. Access devices that may be used to gain minimallyinvasive access to the spine in several of the foregoing surgicalapproaches to the spine include expanding cannula structures such asDubrul U.S. Pat. Nos. 5,183,464 and 5,431,676, Bonutti U.S. Pat. Nos.5,674,240 and 5,320,611, and Davison U.S. Pat. Nos. 6,652,553 and6,187,000.

Laparoscopic Approach

It is also known to approach the lumbar spine anteriorly using alaparoscopic approach. See, for example, Green U.S. Pat. Nos. 5,755,732and 5,620.458. Techniques for laparoscopic placement of spinal fusioncages are shown and described in Kuslich U.S. Pat. No. 5,700,291 andCastro U.S. Pat. No. 6,004,326. Implementing the laparoscopic approachrequires that one or more laparoscopic access devices, commonly referredto as trocars (see for example Moll U.S. Pat. Nos. 4,601,710 and4,654,030) are introduced into the abdominal cavity and that the cavityis insufflated to create working space. A laparoscope is insertedthrough one of the trocar ports to provide visualization of theabdominal cavity and surgical instruments may be introduced eitherthrough another trocar port or through a working channel of thelaparoscope to dissect, manipulate and retract tissue to gain access tothe posterior wall of the abdomen adjacent to the spine. Retractors,including balloon retractors, may be used to retract organs and tissueto maintain a clear working path. Care is taken to avoid damage to themajor blood vessels, the aorta and femoral arteries, and the posteriorwall of the peritoneum is opened to access the desired spinal vertebralbody or disc segment.

C. Imaging Techniques

A variety of tools exist to assist the surgeon in assuring the desiredaccess and treatment are achieved without compromising or adverselyaffecting adjacent healthy tissue. Treatment of the spine is usuallyplanned based on CT or MR scans and fluoroscopy is commonly used duringsurgery to assure proper positioning and placement of surgical tools anddevices. Image guided spinal surgery has been proposed and iscontemplated for use with the surgical treatments proposed herein. See,for example, Cosman U.S. Pat. Nos. 5,662,111; 5,848,967; 6,275,725;6,351,661; 6,006,126; 6,405,072; Bucholz U.S. Pat. Nos. 5,871,445;5,891,034; 5,851,183; and Heilbrun U.S. Pat. Nos. 5,836,954 and5,603,318. The position of instruments typically is detected using acamera and markers on the surgical tool, and an image of the workingportion of the instrument is super-imposed upon a pre-operative image,such as a CT, MRI or ultrasound image to show the surgeon where theworking instrument is located relative to anatomical landmarks and thetissue to be treated. As imaging techniques and equipment improve, it iscontemplated that image guided surgery will evolve to using real timeintraoperative images and that the position of the surgical instrumentwill be shown relative to these real-time intra-operative images inaddition to or in place of pre-operative images.

D. Adhesives and Other Repair Materials.

Numerous patents describe previous approaches to disc repair. Theseinclude U.S. Pat. No. 6,332,894, Stalcup et al., which describes anorthopedic implant for implanting between adjacent vertebrae comprisingan annular bag and a curable polymer and hard particulate with the bag.The polymer is cured after implantation to make it harder and to fusethe hard particulate into a single mass. U.S. Pat. No. 6,264,659, Rosset al., describes a process of injecting a thermoplastic material withinan annulus fibrosis of a selected intervertebral disk. U.S. Pat. No.6,127,597, Beyar et al., describes a solid phase formation device fororthopedic application. The expandable device includes a material thatpolymerizes after implantation. U.S. Pat. No. 6,419,706, Graf, describesa disc prosthesis comprising a preformed polymer core surrounded by arigid material coating. U.S. Pat. No. 6,569,442, Gan et al., describespolymer foam prepared outside the body for intervertebral discreformation.

U.S. Pat. No. 6,022,376, Assell et al., describes a capsule-shapedprosthetic spinal disc nucleus for implantation into a human intradiscalspace, made of a substantially inelastic constraining jacket surroundinga pre-formed amorphous polymer core. U.S. Pat. No. 6,132,465, Ray etal., describes a device similar to the device described in U.S. Pat. No.6,022,376 with certain shape modifications. U.S. Pat. No. 6,306,177,Felt et al., describes an in situ polymerizing fluid used in tissuerepair in the absence of a constraining structure, such as a balloon.The polymerizing materials comprise a quasi-prepolymer component and acurative component containing chain extenders, catalysts and the like.U.S. Pat. No. 4,743,632, Marinovic, discloses the use of a two-partadhesive for use in surgery, where a diisocyanate material is mixed witha polyamine or similar material to produce an in situ cure. Exemplarymaterials are described in U.S. Pat. No. 6,254,327, and our pendingapplications US 2003-0135238 and US-2004-0068078, the disclosures ofwhich are incorporated herein by reference.

E. Other References Providing Background Information

These include U.S. Pat. Nos. Re. 33,258 (Onik et al.), 4,573,448(Kambin), 5,192,326 (Bao et al.), 5,195,541 (Obenchain), 5,197,971(Bonutti), 5,285,795 (Ryan et al.), 5,313,962 (Obenchain), 5,514,153(Bonutti), 5,697,889 (Slotman et al.), 5,755,732 (Green et al.),5,772,661 (Michelson), 5,824,093 (Ray et al.), 5,928,242 (Kuslich etal.), 6,004,326 (Castro et al.), 6,187,048 (Milner et al.), 6,226,548(Foley et al.), 6,416,465 (Brau), WO 01/32100, and FR 2 639 823.

Disclosed herein is the creation of disc nucleus replacement prostheticsthat effectively transfer load on the nucleus, in the form of pressureon the nucleus, to load on the annulus fibrosis in the form of hoopstress. This results in a reduction in nucleus pressure, increase indisc height, and places the elements of the disc in a more normalstructural and load-bearing relationship. Certain ranges of materialsproperties are described that can lead to successful nucleusreplacements, depending on the size of the replacement, the condition ofthe annulus, and the disc height. The insights disclosed herein areequally applicable to prosthetics formed outside the body andprosthetics formed inside the body.

The mechanical properties of the nucleus have been modeled in order todetermine what sorts of materials we should use as nucleus replacements.The modeling involves making assumptions, but has produced useablepredictions for improved materials.

General Considerations for Nucleus Replacement Prosthetics

We start by considering incompressible nucleus replacement prosthetics.By incompressible we mean the implant does not change volume when aforce is applied. Referring to FIG. 1, an incompressible substance 101is loaded along its z-axis 106 as depicted by force vector 102. Theresult is that the solid increases dimensionally along lines 103 in thedirections of the x-axis 104 and y-axis 105. The condition ofincompressibility is given by

V=K or ∀V=0,

where V is the volume of the substance, K is a constant, ∀ is thedivergence operator, yielding the divergence of V.

Accordingly, the effect of stress induced by load 102 on a small unitvolume within the nucleus replacement prosthetic 101 can be seen bytaking the divergence of the product of unit vectors x, y, and z. Sinceload is in one direction, and the substance is isotropic, x isequivalent to y and we need only take the divergence in one direction x.

d/dz(zy ² =V)→y ²+2zydy/dz=0→y=−2dyz/dz→dy/y/dz/z=−½

The last expression is the Poisson ratio, v, where

ε_(z) =dz/z and ε_(y) =dy/y and v=−ε _(y)/ε_(z)

and ε_(y) and ε_(z) are the dimensionless strains in the transverse 103and longitudinal 102 directions, respectively. What these relationsdemonstrate is that a prosthetic that is incompressible under normalphysiologic loads must have a Poisson ratio near ½; and for everypercent decrease in prosthetic height along z 106 there's a ½ percentincrease in prosthetic thickness along axes x 104 and y 105. The ratioof strains in the lateral and transverse directions holds forhomogeneous compressible prosthetics as well, but v<0.5. This means forcompressible compared to incompressible prosthetics there is morestrain, or decrease in height, in the direction of load for the samestrain, or increase in width, perpendicular to the direction of load. Itwill be shown later that compressibility (v<0.5) increases the fractionof the total load at the center of the endplates of the disc.

Disc Kinematics

The natural, healthy spinal disc is comprised of bony endplates defininga disc height, the edges of which are sealed with an annulus bridgingthe endplates and defining a nuclear space. The nucleus is structured,but comparing its modulus to that of the annulus, its modulus isessentially zero. The capacity of the nucleus as an energy storagemechanism is essentially zero and stress applied in the axial directionis immediately translated to stress in the lateral directions, as if itwere a liquid, i.e., σ_(x)=σ_(y)=σ_(z). Solids, on the other hand,resist shape change and for incompressible solids 2σ_(x)=2σ_(y)=σ_(z)for forces applied in the z-direction. This difference means solidsstore energy in compression, liquids do not.

When a load is applied to a disc, work is done in the form ofLε_(z)D_(disc) (Force×Distance) and this energy is stored in the disc aspotential energy. Since the natural nucleus has no energy storingcapability, forces must all be transferred and stored in the annulus.Define the potential (stored) energy in the annulus asE_(p)=E_(hoop)+E_(comp)., where E_(hoop) is the energy stored as hoopstress and E_(comp) is the energy stored as compression stress.

The first mechanism for storing energy in the annulus is hoop stress.Any unconstrained, homogeneous object with internal pressure assumes theshape of a sphere, or if constrained as in the case of the annulus,circular in cross section. Because the annulus is not circular in crosssection, but flattened where it contacts the endplates, the degree offlattening or bowing out of the walls represents stored energy in theform of E_(hoop).

The second mechanism for storing energy in the annulus originates in thecross hatched orientation of the fibers in the annulus. These fibers arenormally at 30 degree angles with respect to the plane of the disc andare interconnected by elastic Type I collagen. Compressive forces reducethe inter-fiber angle and stretch the inter-connecting collagen fibers.The consequence is store energy in the form of Hooke's Law, E=½kε_(z) ²,where k is the spring constant of the fibers.

Given the above conditions of a healthy natural disc, it appears fromthe analysis that any such disc whose nucleus has been replaced by amaterial with E>0, will be stiffer (N/mm) than a natural disc. Sincesuch materials are typical of materials generally used in discreplacement, it is likely that studies that report reduced or equalstiffness in a prosthetic-implanted disc have either damaged one of theenergy storage mechanisms of the disc or have left a compressible voidin the nuclear space. Our analysis indicates that, in contrast toreports in the literature, a nucleus replacement device having with E>0(solid) is not ideal for a healthy, normal annulus. Accordingly,literature studies reporting ideal nucleus prosthetic moduli E in therange of approximately 100 kPa up to the annulus modulus, based onreplacing the nucleus of normal, healthy discs, may be reportingexperimental artifact. Similarly, finite element models found in theliterature that yield such results for a normal, healthy disc arebelieved to reflect computational errors, perhaps introduced by adifferent choice of boundary conditions than hat used in the presentanalysis.

A justification for using a prosthetic with E>0 must presuppose that theenergy storage capacity of E_(hoop) or E_(comp), is reduced from normal,i.e., damaged. To see this, consider a model containing a prostheticwith E>0, resulting in direct energy storage in the prosthetic by anamount Eε_(z) ²A, where A is the area of the prosthetic in contact withthe endplates.

The total energy, Load times Compression, was formerly stored asE_(p)=E_(hoop)+E_(comp.)=constant, in the case where E=0 for thenucleus. Hence, with an E>0 prosthetic, some of E_(p) is stored in theprosthetic as Eε_(z) ²A, reducing E_(hoop) and/or E_(comp). SinceE_(hoop) is rather robust, a likely explanation is that nearly all theloss in the energy storage capacity of the pathological disc is due to abreakdown in the inter-layer collagen connections in the annulusresponsible for E_(comp).

In the theoretical treatment that follows, it is assumed E_(comp)=0.This provides the justification for nonzero prosthetic modulus. Thephysical properties for prosthetics designed to replace E_(comp) overthe entire spectrum of natural disc geometries are derived. Thesesolutions are uniquely obtained by matching E's required to replace lostenergy storage capacity in the natural yet pathologic disc with E'srequired to prevent prosthetic extrusion through the annulotomy made inthe disc during prosthetic insertion. In addition, composite structuresare considered as well as the importance of fixating the prosthetic inthe nuclear space.

One consideration of disc annulus kinematics involves balancing the hoopstresses in the annulus. It will be shown that a consequence ofbalancing the hoop stresses in the annulus is that a unique ideal discdeflection ε_(z) is obtained. Referring now to FIGS. 2A and 2B, the discis shown in lateral cross section 300 and transverse cross section 301.The outer layers of the annulus experience two hoop stresses; σ_(θ) 351and σ_(φ) 352 associated with r_(d) 350, and r_(a) 353, respectively.The angles θ and φ are the usual orthogonal angles in sphericalcoordinates. Accordingly the hoop stresses are given by

σ_(θ) =P _(a) r _(d)/τ and σ_(φ) =P _(a) r _(d)/τ

To balance these stresses, consider that annulus tissue can support morestress in the direction θ, than in φ. It will be shown that thisdifference, while useful in preventing the annulus from rupturing, isrelatively unimportant in determining the ideal disc deflection.Nevertheless, measurements performed on swine annulus indicate that theratio of tensile strength Tθ/Tφ=2.

Therefore, the hoop stresses are balanced when σ_(θ)=2σ_(φ), which givesr_(a)=½r_(d) or more generally r_(a)=(Tφ/Tθ)r_(d)

Referring now to FIG. 3, the annulus in transverse cross section hasdisc height D′_(disc) 361 under mean load and radius of curvature r_(a)362. A bisector x 363 passing through the origin 364 intersects 361 at aright angle and forms angle φ 365.

It follows from trigonometry that

Sin φ=D′ _(disc)/2r _(a)

Now, D_(disc) (loaded) is related to the disc height D_(disc) (unload)by

D _(disc)(unload)=κD _(disc)(loaded)

where κ is the stretch factor or spring constant under mean load.Measurements on swine annulus give a value of κ=0.6.

Then it further follows from trigonometry that

(2φ/π)×2πr _(a) =D _(disc), φ in radians.

Combining these equations yields

2 Sin⁻¹(D′ _(disc)/2r _(a))×2r _(a) =D _(disc)(loaded).

Now D′_(disc)=D_(disc)(1−ε_(z)) and r_(a)=(Tφ/Tθ)r_(d), D_(disc)(unload)=κD_(disc)(loaded), substituting yields

2 Sin⁻¹(D _(disc)(1−ε_(z))/2(Tφ/Tθ)r _(d))×2(Tφ/Tθ)r _(d) =D _(disc)/κ

or

Sin⁻¹(D _(disc)(1−ε_(z))/2(Tφ/Tθ)r _(d))=D _(disc)/(4κ(Tφ/Tθ))r _(d))

which gives

Sin(D _(disc)(4κ(Tφ/Tθ)r _(d))=D _(disc)(1−ε_(z))/2(Tφ/Tθ)r _(d)

Now substituting D_(disc)=9 mm, Tφ/Tθ=½, r_(d)=14 mm gives

Sin(9/28κ)=9(1−ε_(z))/14

For κ=0.60, it follows that ε_(z)=0.21. A disc deflection of 21% undermean load will exactly satisfy σ_(θ)=2σ_(φ).

To show the ideal disc deflection is insensitive to Tθ/Tφ, let Tθ/Tφrange from 1 to 5, which is equivalent to letting σ_(θ)=Sσ_(φ), where Sruns from 1 to 5.

FIG. 4 illustrates insensitivity of the ideal disc deflection under loadfor a large range of ratios of hoop stress, (Tθ/Tφ) for κ=0.6.

FIG. 5 illustrates the sensitivity of the ideal loaded disc deflectionas a function of a narrow range of spring constants of the annulus, κ,for (Tθ/Tφ)=2. It is interesting to note that for an inelastic annulus,(Tθ/Tφ)=2, the ideal loaded deflection is 50%.

In conclusion, an analysis of the spinal disc purely from the point ofview of energy storage and stress balance yields the followingconclusions:

-   -   1. The justification for the use of a prosthetic nucleus with        E>0 is when the energy storage capacity of the annulus has been        impaired. Where the potential energy stored as compression        energy is a minimum and substantially all the stored energy in        the annulus is due to hoop stress.    -   2. Under this condition it can be useful to balance hoop stress        in the disc plane and hoop stress in the transverse plane of the        annulus. This balance yields an ideal loaded deflection of the        disc of approximately 20%.

These findings will be applied in the calculations that follow.

The Disc without a Nucleus

Referring now to FIG. 6, the spinal disc is comprised of bony endplates107, annulus fibrosis 108 and nucleus 109. The annulus is comprised ofrings of fibrous tissue concentric with the center of the disc. Thetensile strength of these rings drops off as their radii decrease. Theoptimal configuration would be a drop off in annular strength of thelayers of the annulus fibrosis consistent with the hoop stress given by

σ_(hoop) =Pr/τ

where σ_(hoop) is the hoop stress, P is the nuclear pressure, r is theradius from the center of the disc to the relevant fibrous layer of theannulus, and τ is the thickness of that layer. This configurationminimizes the shear stress between layers of the annulus fibrosis, whichis a condition for optimal mechanical stability of the disc.

When the inner layers of the annulus fracture or lose their collagencontent, they no longer develop hoop stress and become essentially anextension of the incompressible nucleus 109. Therefore, inner layer 110develops essentially no hoop stress and must be supported by pressurefrom the nucleus. The outer layer 111 develops essentially all the hoopstress that is responsible for the pressure in the annulus 108. It willbe shown below that there are two sources for pressure in the annulus,and they add together to support the overall disc height.

Referring now to FIGS. 7A and 7B, the annulus 108 approximates a torus(as shown in FIG. 7A) with Poisson ratio 0.5 placed between twoendplates 107 (as shown in FIG. 7B) that creates a flat portion 116 anda bulging portion with a radius of curvature 112. Referring to FIG. 7A,the radius of the disc is 113, the radius of the nucleus is 114 and aload placed along the z-axis 106 creates stress in the annulusperpendicular to 112 (hoop stress) and 115 (longitudinal stress).Further, it is recognized that the disc height is approximately twicethe annulus radius of curvature 112. When the nucleus is removed we havethe following static equations of state:

P _(a)=σ_(hoop) τ/r _(a)

P _(a) =L/[π(r _(d) ² −r _(n) ²)]

where P_(a) is the pressure in the annulus, σ_(hoop) is the hoop stressin the annulus, τ is the thickness of the outer annulus layer 111, r_(a)is the radius of the curvature of the annulus 112, L is the load, r_(d)is the radius of the disc 113, r_(n) is the radius of the nucleus 114.

The variables r_(n) and r_(d) will be used to calculate areas over whichloads will be distributed. However, the shape of the disc in lateralcross section, as shown for example in FIG. 7A, is not a circle, butrather a bent ellipse. So, for purposes of application, one needs torelate the surgically measurable quantities, the minor axes of the innerand outer ellipses, r_(n,surg) and r_(d,surg), respectively to theeffective radii.

r_(n)=1.5r_(n,surg) r_(d)=1.5r_(d,surg)

Accordingly, the measured quantities should be increased by 50% fordetermining optimal prosthetic parameters, based on the findingsdisclosed here. For the case of the normal, healthy disc the nuclearvolume equals the annulus volume. This determines the ratio of r_(n) tor_(d)

r_(d)=1.4r_(n) or r_(n)=0.7r_(d)

This ratio determines normal, healthy load sharing between annulus andnucleus. Without nucleus, the disc height, h_(d), can be expressed as afunction of the other parameters.

σ_(hoop) τ/r _(a) =L/[π(r _(d) ² −r _(n) ²)]→r _(a)=σ_(hoop)τ[π(r _(d) ²−r _(n) ²)]/L→h _(d)=σ_(hoop)τπ(r _(d) ² −r _(n) ²)/L

It can be seen that as load, L, increases the disc height decreases. Theeffect of decreasing the annulus width (increasing the nucleus width),which can occur during nuclectomy, can be dramatic in a disc withoutnucleus, as shown below.

FIG. 8 is a plot of the disk height (y-axis) at constant load as thenucleus radius (x-axis, % total disc) is increased. In FIG. 8, the discheight is normalized to 1 when the annulus spans the entire diskdiameter yielding the maximum load bearing capacity. The nucleusthickness or radius is r_(n)=1 when the annulus width is 0(r_(d)−r_(n)).

The Disc with an Incompressible, Zero Modulus Nucleus

Now let's derive the same relation when the disc is filled with anatural nucleus, with Poisson ratio 0.5, and Modulus 0 (fluid). Themodulus 0 condition is valid under static loads, since the nucleus willalways flow to make stress in the z-direction 106 zero. Or statedanother way, the pressure or stress in the nucleus is everywhere thesame when constrained, and zero when not constrained. Then we have thefollowing state equations:

L_(n)=πr_(n) ²P_(n)

L _(a)=[π(r _(d) ² −r _(n) ²)]P _(a)

L=L _(a) +L _(n)

P _(a)=σ_(hoop) τ/r _(a)

where L_(n) is the load supported by the nucleus, and L_(a) is the loadsupported by the annulus. Then summing gives the total load, L.

L=τr _(n) ² P _(n)+[π(r _(d) ² −r _(n) ²)]P _(a)

Now for nuclear modulus 0, P_(a)=P_(n), then

L=π(r _(d) ²)σ_(hoop) τ/r _(a)

which gives

h _(d)=σ_(hoop)τπ(r _(d) ²)/L disc with nucleus

h _(d)=σ_(hoop)τπ(r _(d) ² −r _(n) ²)/L disc without nucleus

FIG. 9 is a plot of % disc height lost (y-axis) when the nucleus isremoved, as indicated by nucleus to disc radius, r_(n)/r_(d) (x-axis).FIG. 9 illustrates why replacing the nucleus after nuclectomy isperformed.

Thus, we see the disc height for disc with-nucleus is greater than thedisc height for disc without-nucleus for all loads L. The relativecontributions, R_(h), of the annulus and nucleus to supporting discheight are given by

$\begin{matrix}{R_{h} = {{height}\mspace{14mu} {contribution}\mspace{14mu} {from}\mspace{14mu} {{nucleus}/}}} \\{{{height}\mspace{14mu} {contribution}{\mspace{11mu} \;}{from}\mspace{14mu} {annulus}}} \\{=} \\{= {{\left( {h_{n} - h_{a}} \right)/h_{a}} = {r_{n}^{2}/\left( {r_{d}^{2} - r_{n}^{2}} \right)}}}\end{matrix}$

where h_(n) is the disc height for a disc with nucleus, h_(a) is theheight contribution from the annulus alone. For normal discs, r_(d)≅1.4r_(n), so

R_(h)=1

This expression illustrates the importance of filling the nucleuspost-nuclectomy, and that filling an empty nucleus with anincompressible fluid (E=0) improves the disc height by 100% under allphysiologic loading conditions.

Returning to our previous equations:

L_(n)=πr_(n) ²P_(n)

L _(a)=[π(r _(d) ² −r _(n) ²)]P _(a)

The load ratio, R_(I), is given by

R _(I) =L _(n) /L _(a) =r _(n) ²/(r _(d) ² −r _(n) ²)=1

when r_(n)/r_(d)=0.7

FIG. 10 illustrates that load is distributed between the annulus andnucleus according to the ratio of the radii of the nucleus and disc.FIG. 11 shows the normal load distribution in an isolated lumbar (L5)vertebral segment from a 200 lb pig. This data validates the relationr_(d)=1.4r_(n).

As expected, for a healthy disc r_(n)/r_(d)=0.7, the loads aredistributed as a linear superposition, with about 50% of the loadsupported by the annulus alone. It will be demonstrated below that theannulus contribution to the load support drops as the modulus of thenucleus increases.

The Disc with Incompressible Nucleus Having Modulus Greater than Zero

The pressure ratio R_(p) for an E=0 nucleus, as stated before asP_(a)=P_(n), is 1. The pressure ratio R_(p) will be greater than 1 whena nucleus with a non-zero modulus, E, is used to replace the naturalnucleus. In other words, as the modulus of the nucleus increases, moreof the load is transmitted to the soft center of the endplates where thenucleus is positioned with less being transmitted to the harderperiphery of the endplates where the annulus is located. To see this,the concept of a modulus must be introduced. Simply stated, the modulusdetermines the distribution of stresses in a solid based on strain, forexample

E=σ _(xx)/ε_(xx)

where E is Young's modulus, σ_(xx) is the stress along the x-axis andε_(xx) is the strain. For convenience, stress and strain will only begiven along the principal axes x, y, z, so the double subscript can bereplaced by a single letter. From the Poisson relation, v=−ε_(x)/ε_(z),we get

v=0.5=−ε_(x)/ε_(z)=−σ_(x)/σ_(z)

So there's twice the amount of stress and strain in the direction ofload than at the bulging edges. In terms of a disc under load L indirection z, ε_(z) is the fractional disc height shortening, Δ(discheight)/(disc height), called here disc deflection. Therefore, we havethe following state equations:

L _(n) =πr _(n) ²(Eε _(z) +P _(a))

L _(a)=π(r _(d) ² −r _(n) ²)P _(a)

L=L _(a) +L _(n)

Note that when E=0 the state equations for the natural nucleus areretrieved. If Eε_(z) and P_(a) are left uncoupled, these equations havesolutions that place unsafe stress on the endplates of the vertebralbodies. To prevent this from happening, a boundary condition is applied,called here the Endplate Limit. It is a direct consequence of fillingthe nucleus to capacity such that the inner layers of the annulus andthe prosthetic are in equilibrium. When this condition is fulfilled,P_(a)=Eε_(x). This is a rigid requirement, but there is a physiologicalreason as well.

In FIG. 12, a disc 250 is shown in transverse cross section withendplates 107 and annulus 108. Prosthetic 251 resides in contact withannulus 108. To define the Endplate Limit, consider first the geometryof the annulus. Depicted in out-take 252 is the layer orientation of theannulus. The endplates are made up of a layer of fibrous tissue that isessentially a continuation of the annulus, which forms a type ofenveloping sack around the nucleus. Underneath this layer is bone. Theorientation of the fibrous layers on the endplates is depicted in outtake 253. Together, the tissues depicted in 252 and 253 form acontinuous structure, the layers of which are approximately orientedperpendicular to a vector with origin at the center of the nucleus. Itis this direction 254 for the annulus 108 and 255 for the endplates thatprovides the most resistance to nuclear pressure. These layers areweakest in the in-plane directions 255 and 256. In the case of theannulus, the force aligned in the in-plane direction is the hoop stressσ_(hoop)=P_(a)r_(a)/τ, where P_(a) is the annulus pressure, r_(a) is theradius of curvature, and τ is the wall thickness. In the case of theendplates, the in-plane stress due to nuclear pressure iscounter-balanced by adhesion to the adjacent bony layer of the vertebralbody. However, with the addition of prosthetic 251, there is anadditional stress Eε_(x), and it is this stress in the weak in-planedirection, and not Eε_(z) in the strong direction 255 that requireslimitation. The Endplate Limit essentially establishes a balance betweenforces that promote endplate failure, Eε_(x), and forces that promoteannulus failure, σ_(hoop). Mathematically, this is expressed as

σ_(hoop)=Eε_(x).

The annulus is essentially a solid inner tube, where the wall thicknessequals the radius of curvature, r_(a)=τ. Substituting this relation intothe equation σ_(hoop)=P_(a)r_(a)/τ=Eε_(x) gives the Endplate Limit

P_(a)=Eε_(x) Endplate Limit

Now, Eε_(z)=2Eε_(x)., for v=0.5, and Eε_(x)=Eε_(y), substituting theseexpressions we arrive at an alternative form of the Endplate Limit

Eε _(z) =Eε _(x) +P _(a)

The Endplate Limit has a rather elegant consequence that for the normalnucleus case r_(d)=1.4r_(n), the load on the annulus is reduced by 50%,from 50% of total load to 25%. This ratio, 25% for the annulus and 75%for the nucleus, is constant for all disc deflections, as will be shownbelow using the following equation derived from the state equations.

L _(n) /L _(a)=(r _(n) ² /r _(d) ² −r _(n) ²){ε_(z)(E,L _(n))/[ε_(z)(E,L_(n))−ε_(x)(E,L _(n))]+1}.

FIG. 13 is a plot showing load on the nucleus/load on annulus(L_(n)/L_(a), y-axis) versus disc deflection (x-axis). Comparing FIG. 10to FIG. 13, replacing the natural nucleus (E=0) with a nucleusprosthetic (E>0) shifts load to the nucleus. This is a practicalnecessity in cases where the annulus has already demonstratedinsufficiency under load. For the case r_(n)/r_(d)=0.7, the load ratioL_(n)/L_(a) shifts from 1 (E=0) to 3 (E>0). Note that by obeying theEndplate Limit constraint the load ratio is a function of r_(n)/r_(d)and not the prosthetic modulus or the load.

The disc deflection ε_(z) is dependent on the modulus E and the load Lon the total disc. It is useful to separate the nucleus load componentsin terms of nuclear pressure L_(p) (P_(n)=P_(a)) and the load supportedby the modulus of the prosthetic L_(m).

L_(m) = π r_(n)²E ɛ_(z) L_(p) = π r_(n)²P_(a) Then$\begin{matrix}{{L_{m}/L} = {\pi \; r_{n}^{2}{ɛ_{z}/\left\{ {{\pi \; {r_{d}^{2}\left\lbrack {{ɛ_{z}\left( {E,L_{n}} \right)} - {ɛ_{x}\left( {E,L_{n}} \right)}} \right\rbrack}} + {\pi \; r_{n}^{2}ɛ_{z}}} \right\}}}} \\{= {r_{n}^{2}{ɛ_{z}/\left\lbrack {{\left( {r_{n}^{2} + r_{d}^{2}} \right)ɛ_{z}} - {r_{d}^{2}ɛ_{x}}} \right\rbrack}}} \\{= R_{m}}\end{matrix}$

The fraction of the entire load L supported by the prosthetic, andtherefore also by the soft centers of the endplates is explicitlydetermined by the ratios of r_(n) and r_(d). This is a clinically usefulresult since the surgeon can match the modulus of the implant based onthe amount of viable annulus remaining and the total deflection ε_(z)desired under a given load L. The Young's modulus is then given by

E=LR _(m) /πr _(n) ²ε_(z)

For 1.4r_(n)=r_(d), then R_(m)=0.505 and E=0.161 L/r_(n) ²ε_(z). E is inmegapascals if r_(n) is in mm and L is in Newtons.

The correct modulus for a disc prosthetic can be determined byconsidering the normal Load vs. Displacement relation of a disc, asillustrated in FIG. 14, which shows the displacement (x-axis) as afunction of load (y-axis). When compressing a disc by amountD_(disc)ε_(z) for a given nucleus modulus, L is proportional toε_(z)/(1−ε_(Z)), as depicted in FIG. 14, where typically L=1000 N whenD_(disc)ε_(z)=1 mm and ε_(z)=20%.

However, the equation above gives L proportional to ε_(z) for fixed E.This is a direct consequence of incorporating the Endplate Limit intothe equation for selecting appropriate modulus E. Accordingly, for anytarget pair (ε_(z),L) as shown in FIG. 14 by the dashed lines, aprosthetic modulus E is specified by the equation that satisfies theEndplate Limit.

Considering a collection of points (ε_(z),L) on the line graphed in FIG.14, for each point there is a unique and different prosthetic modulus Eproportional to (1−ε_(z))⁻¹, that balances the loads such that theEndplate Limit is satisfied. For a trace such as given in FIG. 14 thereis only one value E corresponding to each paired value (ε_(z),L) forwhich the Endplate Limit is satisfied. When the clinician selects E forvalues of L=1000 N and ε_(z)=0.20 (1 mm) he is selecting a modulus forthe prosthetic which will result in the implanted disc reproducing thisnormal disc relation between L and ε_(z). These are the load anddisplacement values that are typical for the loaded spine.

For discs ranging from normal to pathologic, r_(n)>=0.7 r_(d), there isno non-zero value for E which gives R_(L)=1 for any values of L andε_(z). The condition R_(L)=1 is the loading condition for the healthydisc at r_(n)=0.7r_(d). The only way to achieve R<3 for r_(n)>=0.7 r_(d)is to use a liquid/solid composite prosthetic. This case will be coveredin the section entitled, the Disc with Bubble Entrapped Nucleus.

FIGS. 15-17 illustrate additional consequences of the EndplateLimitation restriction.

FIG. 15A illustrates that for fixed r_(n)/r_(d)=0.7, the absolute sizeof the nucleus determines implant modulus when the desired deflectionand load are specified.

FIG. 15B illustrates the requirements on implant modulus when the ratior_(n)/r_(d) is not fixed, but r_(d)=20 mm is fixed. The modulusrequirement drops with increasing implant radius, r_(n).

FIG. 16 illustrates the modulus required for the normal disc conditionr_(d)=1.4r_(n) as a function of anticipated maximum load. This figuregenerally illustrates that for fixed deflection, and r_(n)/r_(d)=0.7,the modulus requirement increases linearly with increasing load.

FIG. 17 illustrates the sensitivity of the modulus requirement on thetarget deflection under a fixed load of 1000 N. A disc height of 5 mmhas been used in these calculations to represent the state of apathological disc. Normally, in the lumbar region disc height is between9 and 12 mm. Restoring a disc to this height by some means ofdistraction will afford a larger range of deflection under load. Forexample, a surgeon may elect to choose a prosthetic moduluscorresponding to deflections of 40-50%. The above FIG. 17 illustratesthat doing so will halve the required modulus.

The target deflection under load will be determined to some extent byannulus health. Annulus health is reflected in the ratio r_(n)/r_(d),which has been taken to be a healthy 0.7 in the example above. Althoughincreasing r_(n) by removing annulus decreases the requirements on themodulus, the ratio R_(L) increases as r_(n)/r_(d) increases. This isshown in FIG. 13. These illustrations serve to emphasize the importanceof surgically preserving as much of the viable annulus as possible toobtain the lowest possible value of r_(n)/r_(d) and R_(L). The surgeonwill likely find FIG. 13, where r_(n)/r_(d) varies, to be most usefulenabling him to select the ratio consistent with the outcome of thenuclectomy.

In the limit, E=0, the natural disc places no uncompensated in-planestress on the endplates, and the entire load on the nucleus is carriedin the z-direction of the endplates. Disclosed herein are therequirements of a failing annulus. The choice was made to transfer someof the stress normally in the annulus to in-plane stress in theendplates. Utilizing this aspect of the endplates provides support fordisc height that is analogous to hoop stress, without placing thatstress in the annulus. This is achieved by filling the nuclear space tocapacity such that the prosthetic is in pressure equilibrium with theannulus. Strategies that do not establish equilibrium with the annulusplace 100% of the restorative force generated by the prosthetic in thez-direction on the endplates, and concentrate those forces in an areasmaller than the cross sectional area of the nucleus.

It should be recognized that the Endplate Limits established in thepreceding are for disc heights of 5 mm. This is a typical height for adiseased disc. Therapies that do not increase disc height must providefor these high moduli. It is one object of these analyses to demonstrateand emphasize the importance of combining a nucleus replacement therapywith either disc height augmentation or prosthetic bonding to theendplates. When disc height augmentation is performed, prosthetic moduliare typically reduced by at least 50%. When prosthetic bonding isachieved even greater reductions in prosthetic modulus are possible.

The Disc with a Nucleus Prosthetic Bonded to its Endplates

The foregoing assumed the nucleus, natural or replaced, is homogenous inmodulus. Accordingly, the stress in the nucleus is isotropic in a planetransverse to the direction of the load σ_(x)=σ_(y) for all planes in z.Due to cyclic boundary conditions imposed by the toroidal geometry ofthe annulus, it follows ε_(x)=ε_(y) for all planes in z. In the casewhere the prosthetic is bonded to the endplates, the transverse strainvaries as a second order polynomial of z, and the effective modulus isno longer homogenous in z.

Under the condition σ_(x)<bond strength in shear, Hooke's law may beapplied, where the elastic modulus tensor C_(i,j,k,l) is summed over kand l. However, the bond at the endplates yields this simplified form ofthe stress

σ_(ij)=C_(i,j,k,l)ε_(kl) =C ₁₁₁₁ε₁₁ +C ₁₁₂₂ε₂₂ +C ₁₁₃₃ε₃₃ =C ₁₁₁₁ε₁₁

so the effective stiffness for constrained compression is C₁₁₁₁.

Solving for C₁₁₁₁=σ_(z)/ε_(z), which is the modified modulus E, in termsof E starts with the elementary form of the isotropic Hooke's law:

ε_(xx)=(1/E){σ_(xx) −vσ _(yy) −vσ _(zz)}

ε_(yy)=(1/E){σ_(yy) −vσ _(xx) −vσ _(zz)}

ε_(zz)=(1/E){σ_(zz) −vσ _(xx) −vσ _(yy)}

Then impose the constraint ε_(xx)=ε_(yy)=0, then

σ_(yy) =vσ _(xx) −vσ _(zz)

σ_(xx) =vσ _(yy) −vσ _(zz)

Substituting

σ_(yy)=σ_(xx)=σ_(zz) {v(1+v)/(1+v)(1−2v)}

Then

C ₁₁₁₁=σ_(z)/ε_(z) =E{(1−v)/(1+v)(1−2v)}

Here, v=0 at the bonded ends and approximates v=r_(n)/D_(disc) whereD_(disc) is the disc height or prosthetic height in the direction z. Andthe equations of state become

L _(n) =πr _(n) ² [Eε _(z){(1−v)/(1+v)(1−2v)}+P _(a)]

L _(a)=π(r _(d) ² −r _(n) ²)P _(a)

Eε _(z) =Eε _(x) +P _(a)

And the load ratio R_(I) is

R _(I) =L _(n) /L _(a) =[r _(n) ² /r _(d) ² −r _(n)²]{[(1−v)ε_(z)/(ε_(z)−ε_(x))(1+v)(1−2v)]+1} bound

R _(I)=(r _(n) ² /r _(d) ² −r _(n) ²){ε_(z)(E,L _(n))/[ε_(z)(E,L_(n))−ε_(x)(E,L _(n))]+1}. unbound

This is illustrated in FIG. 18. Here, the Poisson ratio v is theeffective ratio due to prosthetic bonding to the endplates, and isgreater than ½. The result of endplate bonding of the prosthetic is thatmore load is shifted to the endplates when compared to the loaddistribution of the same prosthetic not bound to the endplates. Recallthat decreasing the modulus of the prosthetic shifts load to theannulus. Therefore, when the endplates are bonded a lower modulusprosthetic may be used to achieve the same distribution of loadsachieved for a free prosthetic of higher modulus.

The Disc with Bubble Entrapped Nucleus

Gaseous inclusions in a water permeable prosthetic will exchange gaseswith the surrounding fluids in the tissue. During periods of minimalloading the spring constant of the bubbles will result in their gaseousvolume being replaced with a liquid volume. Let the total volume of theliquid component of the implant be expressed as a fraction of the totalimplant volume, v_(L). Let v_(P) be the non-liquid volumetric fractionof the prosthetic, where v_(L)+v_(P)=1. The equations of state are

L_(L)=πr_(n) ²P_(a)v_(L)

L_(M)=πr_(n) ²Eε_(z)v_(P)

L_(P)=πr_(n) ²P_(a)

L _(a)=π(r _(d)2−r _(n) ²)P _(a)

L _(n) =L _(L) +L _(p)

where L_(L) is the load supported by the liquid in the prosthetic, L_(M)is the load supported by the modulus of the prosthetic, L_(P) is theload supported by the nuclear pressure. Then

L _(n) /L _(a) =r _(n) ²(1+v _(L))/r _(d) ² −r _(n) ² +r _(n) ²ε_(z)(E,L_(n))v _(P)/(r _(d) ² −r _(n) ²)[ε_(z)(E,L _(n))−ε_(x)(E,L _(n))]

where

R _(m) =v _(P)/[(r _(d) ² /r _(n) ²+1)−(r _(d) ² /r _(n) ² +v _(L))]v

E=LR _(m) /πr _(n) ²ε_(z) v _(P)

The effect of the bubbles is to reduce the load on the center of theendplates.

Disc with Partially Filled Nucleus

Disclosed herein are optimal mechanical properties for a nucleusprosthetic intended to entirely replace the natural nucleus. Thefollowing details the disadvantage of partially, rather than completelyfilling the nuclear space.

In earlier attempts to provide a replacement for the nucleus someportion of the natural nucleus was left in the disc. These nucleusreplacements had modulus greater than 0, and consequently, naturalnucleus would extrude around the implant and out a defect in theannulus. In the case where the annulus is sealed, the results for thispartial replacement of the nucleus would be similar to those findingsreported in the section The Disc with Bubble Entrapped Nucleus.

In the case where the entire natural nucleus is removed and only aportion of the nuclear space is filled by a replacement prosthetic and aspace remains, then we have the equations of state:

L_(n)=πr_(p) ²Eε_(z)

L _(a)=π(r _(d) ² −r _(n) ²)P _(a)

where the radius of the prosthetic r_(p)<r_(n). Now, the pressure in theannulus is

P _(a)=(L−πr _(p) ² Eε _(z))/π(r _(d) ² −r _(n) ²) prostheticradius<nuclear radius

Compared with

P _(a) =L−πr _(n) ² Eε _(z)/πr_(d) ² prosthetic radius=nuclear radius

FIG. 19 illustrates r_(d)=14 mm, r_(n)=10 mm, E=3 MPa, L=1000 N,ε_(z)=0.2. FIG. 19 shows that the burden placed on the annulus, P_(a),is much greater when the prosthetic does not fill the nuclear space.This is partially due to the fact that the portion of the load supportedby the prosthetic decreases with smaller prosthetic size (the slopingportion), and partially due to decoupling of the nucleus from theannulus (the offset). Even a prosthetic that fills 90% of the nuclearspace would achieve a further reduction in annulus pressure of about 50%if coupled to the annulus with water.

Nucleus Extrusion Pressure and Prosthetic Failure

It should not be ignored that often discs that undergo nuclectomyrequire an annulotomy or already possess a defect that must be removed.Therefore, any nucleus replacement must be designed to prevent extrusionof the nucleus prosthetic through the annulotomy.

The forces involved in prosthetic extrusion through the annulus can bestbe understood by referring to FIG. 20. Note a labeling change where themain direction of force is the z-axis, which for extrusion in the discannulus is in the former xy-plane. An equilibrium is maintained betweenthe forces acting on the elemental slice 200 of the prosthetic as itextrudes through static material zone 203. The stresses acting onelement 200 are shown in FIGS. 21A-21C. The equilibrium equation isgiven by

−(p _(z) +dp _(z))π(D+dD)²/4+p _(z) πD ²/4+p _(r) πDds sin α+τ_(f) πDdscos α=0

where τ_(f) is the frictional stress in the dead-zone 213, p_(r) is theradial pressure 210, α is the dead-zone angle 214, p_(z) is the pressurein the z-direction 211, p_(z)+dp_(z) is 212, ds is the length 217, D isthe length 215, D+dD is length 216, 202 is the disc height D_(disc), 205is the element width dz, the length z is 204, 201 is the diameter of theannulus defect D_(D).

The equilibrium equation can be further simplified by using thefollowing relationships among dz 205, dD, and ds 217:

ds sin α=dz tan α=dD/2

ds cos α=dz=dD/2 tan α

The von Mises yield criterion is

p _(r) =p _(z)+σ and τ_(f)=σ/3^(1/2)

where σ is the flow stress in the prosthetic. Substituting andneglecting higher order differentials, the equilibrium equation isobtained in the integral form

dp _(z)/[σ(1+cot α/3^(1/2))]=2dD/D

Assuming flow stress is constant during extrusion, the integration ofthe equation yields

log_(e) D ² C=p _(z)/[σ(1+cot α/3^(1/2))]

where C is the integration constant.

The boundary conditions D=D_(D) (diameter of annulus defect) andp_(z)=0, give an expression for C

C=D_(D) ⁻²

Substituting back into the equilibrium equation for the constant C, theaverage extrusion pressure is

P _(ave,z=0)=2σ(1+cot α/3^(1/2))log_(e)(D _(disc) /D _(D))

where D_(disc) is the disc height.

The nucleus/defect interface friction must be included to determine thetotal pressure required for extrusion. The equation for staticequilibrium in the z-direction is

[(p _(z) +dp _(z))−p _(z) ]πD _(disc) ²/4=τ_(f) πD _(disc) dz

where τ_(f) is the frictional force at the nucleus/defect interface. Inthe integral form

dp _(z)/τ_(f)=4dz/D _(disc)

Integrating, and substituting the boundary condition at z=0,p_(z)=P_(ave,z=0), the extrusion pressure is

p _(z)=4τ_(f) z/D _(disc) +P _(ave,z=0)

Now substituting the expression for P_(ave,z=0) and τ_(f) at yield, theaverage extrusion pressure is

P _(ave)=2σ(1+cot α/3^(1/2))log_(e)(D _(disc) /D _(D))+4σz/3^(1/2) D_(disc)

Now σ=Eε where

ε={(12VD _(disc) tan α)(D _(disc) ³ −D _(D) ³)}2 log_(e)(D _(disc) /D_(D))

and ε is the mean velocity strain, V is the impact speed in mm/s of thevertebral load.

Also, from geometric considerations

$\alpha = {\tan^{- 1}\left\lbrack {\frac{1}{2}{\left( {D_{disc} - D_{D}} \right)/\left( {r_{d} - r_{n}} \right)}} \right\rbrack}$$\begin{matrix}{z = {r_{n} - r_{a}}} \\{= {{0\mspace{14mu} {iff}\mspace{14mu} r_{n}}<=r_{a}}}\end{matrix}$

where r_(d) is the disc radius, r_(a) is the annulus radius of curvatureand r_(n) is the nucleus radius.

The explicit expression for the average extrusion pressure is

P _(ave)=2E{{(12VD _(disc) tan α)/(D _(disc) ³ −D _(D) ³)}2 log_(e)(D_(disc) /D _(D))}××{(1+cot α/3^(1/2))log_(e)(D _(disc) /D _(D))+2(r _(n)−r _(a))/3^(1/2) D _(disc)}

It is worthwhile to note that P_(ave) is the nuclear pressure P_(n) inearlier calculations, and not the load. The load required to causeprosthetic extrusion through the annulus is

L=L _(n) +L _(a) =πr _(n) ²(Eε _(z) +P _(ave))+π(r _(d) ² −r _(n) ²)P_(ave)

Let D_(D)=2.5 mm, D_(disc)=5 mm, r_(n)=10 mm, r_(a)=2.5 mm and itfollows that r_(d)=14 mm, z=7.5 mm, tan α=0.3125, cot α=3.2. ThenP_(ave) and the corresponding Load at failure is

P_(ave)=2.21EV unbound

P_(ave)=19.9EV bound

L _(fail)=314Eε _(z)+1360EV unbound

L _(fail)=314Eε _(z)+12,200EV bound

For the “at rest” case, let V=1 mm/s, then the failure pressure P_(ave)for all deflections as a function of the modulus E is illustrated below.

FIG. 22 illustrates that failure pressure rises linearly with prostheticmodulus when everything else is fixed and the implant is not bonded tothe endplates. A high failure pressure is desirable.

FIG. 23 illustrates that the faster the load is applied the higher thefailure pressure. For static load, the disc has a characteristicrelaxation time to reach equilibrium. The relaxation time T_(r) dividedby the change in disc height ε_(z)D_(disc) is the impact velocity V inthe static load case.

V_(static)≅ε_(z)D_(disc)/T_(r)≅1 mm/s

Older discs will typically equilibrate faster, making V_(static) larger.

FIG. 24 is load failure as opposed to nuclear pressure failure,illustrated previously. Fairly large load failures are generated byrelatively modest implant modulus.

Now consider a large annulus defect. Let D_(D)=4.9 mm, D_(disc)=5 mm,r_(n)=10 mm, r_(a)=2.5 mm and it follows that r_(d)=14 mm, z=7.5 mm, tanα=0.0125, cot α=80.

P_(ave)=0.0221EV

L=314Eε _(z)+13.6EV

FIG. 25 illustrates a large decrease in load failure threshold forlarger defect size. Despite the fact that a disc with a large defect(4.9 mm) will have a high V, this consequence does not appreciablyimprove the load failure threshold. On the other hand, for a healthierannulus (defect=2.5 mm) the impact velocity has a large effect on theload failure threshold for near static loading. The effect of prostheticbonding is also illustrated

FIG. 26 provides a convenient conversion from modulus (MPa) to durometer(Shore A) for polyurethane prosthetics.

FIGS. 27 and 28 illustrate load thresholds for free prosthetic forvarious deflections where D_(D)=2.5 mm. These plots describe a cone,within which the spectrum of allowed spinal motion is bound. The conenarrows for increasing V.

Now, consider the following conditions D_(disc)=5 mm, r_(n)=10 mm,r_(a)=2.5 mm, r_(d)=14 mm, deflection=0.20, z=7.5 mm where D_(D)=0.5, 1,2.5, 3.5, 4.9.

Failure Thresholds for Various Defect Sizes

Failure Pressure (MPa) Failure Load (N)

P_(ave,0.5)=15.9EV L=314Eε _(z)+9807EV

P_(ave,1.0)=8.10EV L=314Eε _(z)+4987EV

P_(ave,2.5)=2.21EV L=314Eε _(z)+1360EV

P_(ave,3.5)=0.624EV L=314Eε _(z)+384EV

P_(ave,4.9)=0.0221EV L=314Eε _(z)+13.6EV

P_(ave)≅□□0.84[(Ddisc/D_(D))^(1.3)−1]EV L=314Eε _(z)+264[(Ddisc/D_(D))^(1.3)−1]EV

For r_(d)=1.4r_(n)For bonded

P_(ave) ^(bond)=9.0P_(ave) and L_(bond)=9.0L.

This is illustrated in FIGS. 29A and 29B.

The Disc in Flexion/Extension

To this point the load vector was always perpendicular to the plane ofthe disc. In what follows, we treat the case of flexion-extension underload. Referring to FIG. 31, the midline 220 makes an angle Ω 221 to theplane of the endplates 222. The load L, 223, is a vector withperpendicular components 224, 225. Component 224 is in the plane of theendplate and does not contribute to compression of the disc. Somewhere,at a distance X 231 from the edge of the endplate is a pivot 230coplanar with the endplate. Then the torques 227, 228 and 229 balance.Referring to FIG. 32, the annulus 231 and nucleus 232 apply forces tothe endplate 222 proportional to the area, where φ is the sweeping angle233 defining infinitesimal element 234 and the forces in integral formare:

Annulus: 2(r_(d)−r_(n))P_(a)(φ)sin φdx

Nucleus: 2r_(n)[P_(n)+Eε_(z)(φ)] sin φdx

where x=2r_(d)+X−r_(d)(1+cos φ) and dx=r_(d) sin φdφ

The equilibrium equation then is

Int(φ,0→180){[2r _(d) +X−r _(d)(1+cos φ)]2r _(d) ² sin² φP _(a)dφ}−−Int(φ,0→180){[X+r _(n) +r _(d) −r _(n)(1+cos φ)]2r _(n) ² sin² φP_(a) dφ}++Int(φ,0→180){[X+r _(n) +r _(d) −r _(n)(1+cos φ)]2r _(n) ² sin²φ[P _(n) +Eε _(z)(φ)]dφ}==L cos φ[X+r _(d)/2]

This equation separates into a perpendicular component (α=0) and a tiltcomponent (α>0). Referring to FIG. 33, without tilt the disk heightwould be 240, D_(disk,mean) and with tilt it would increase on theposterior side by 241, r_(d) sin Ω and decrease by the same amount onthe anterior side 242. This is the case because there is no forceexerted at the pivot in the equation above, in fact X can be taken to beany length. Now coupling the mean annulus pressure P_(a,mean) andannulus radius of curvature r_(a,mean) to the mean deflection ε_(z) wehave:

P _(a,mean)=(1/πr _(n) ²)L−Eε _(z)

r _(a,mean)=(D _(disc) −D _(disc)ε_(Z))/2

Then we get for posterior and anterior annulus pressures:

P _(a,post)=(r _(a,mean) /[r _(a,mean) +r _(d) sin Ω])P _(a,mean)

P _(a,ant)=(r _(a,mean) /[r _(a,mean) −r _(d) sin Ω])P _(a,mean)

The nuclear pressure has three components in the tilt case, anequilibrium component equal to P_(a,mean), and two translationalcomponents Eε_(z) sin Ω and (P_(a,ant)−P_(a,post)).

P _(n)={(1/πr _(n) ²)L−Eε _(z) }+Eε _(z) sin Ω+(r _(a,mean) r _(d) sinΩ/[r ² _(a,mean) −r _(d) ² sin²Ω])[(1/πr _(n) ²)L−Eε _(z)]

To avoid extrusion

P_(n)<P_(ave)

Note, the translational component changes the previous physicalinterpretation of P_(n). Now calculate P_(n) for the followingconditions D_(disc)=5 mm, r_(n)=10 mm, r_(d)=14 mm, D_(D)=2.5,deflection=20% and Ω=7 degrees. Than r_(a,mean)=2.0 mm and usingR_(m)=r_(n) ²ε_(z)/[(r_(n) ²+r_(d) ²)ε_(z)−r_(d) ²ε_(x)]=0.5

E=LR _(m) /πr _(n) ²ε_(z)=0.00795L

P _(n)=0.0123L−0.77E=0.776 E tilt

P_(n)=0.177E without tilt

P_(ave,2.5)=2.12E failure threshold, D_(D)=2.5, V=1 mm/s

For the nucleus not to extrude through a 2.5 mm diameter hole in theannulus under 7% flexion, Pn, the internal nuclear pressure must be lessthan the failure threshold for that defect size, P_(ave,2.5). Thefailure pressure is not exceeded, and the implant is not extruded.Therefore, flexion/extension does not exceed the Endplate Limit. Forlesser constraints, the implant is likely to fail under flexion becausethe nuclear pressure increases 438%.

It appears counter intuitive that the nuclear pressure should increasewith E, but one should keep in mind that this equation is valid for onlyone value of deflection=20%, therefore to maintain the same deflectionfor higher E the load must be increased, hence rising nuclear pressure.

The graph of FIG. 30 illustrates that any defect less than approximately3.5 mm in diameter would not fail if a free prosthetic has a modulus of1 MPa or greater and the tilt angle is less than 7 degrees. Withprosthetic bonding, the failure threshold increases to a defect size of4.8 mm in diameter.

The Disc in Torsion

Torsion principally acts to reduce disc height. The height is reduced bytorsion angle Ψ where

D_(disc)=D_(disc) cos Ψ

For angle ψ=12 degrees, D_(disc) is decreased by 2%, a negligibleamount. Torsion does not significantly alter prosthetic moduluscharacteristics derived by applying the Endplate and Extrusion Limits.

Surgical Insights

The following are procedural elements that can be useful in the surgicalreplacement of a natural nucleus with a prosthetic nucleus:

-   -   1. The following conversion factors are to be applied for        converting measured disc radius, r_(d,surg), and nucleus radius,        r_(n,surg), to their effective equivalents used in the formulae        developed here.

r_(n)=1.5r_(n,surg) r_(d)=1.5r_(d,surg)

-   -   2. The Endplate Limit, P_(a)=Eε_(x), is applied to all tables        for selecting ideal prosthetic properties. This condition        ensures that if a prosthetic were placed in a healthy nucleus        that load would be optimally supported by the annulus and the        endplates. It is derived from balancing the in-plane stress in        the annulus with the in-plane stress on the endplates.    -   3. The disc Deflection Limit, ε_(z)=20% at maximum (1000 N)        load, is applied to all tables for selecting ideal prosthetic        properties. This condition ensures that under normal loading        conditions orthogonal hoop stresses in the annulus are balanced        and a disc with its nucleus replaced with a prosthetic deflects        the same amount as a natural healthy disc. Under this        constraint, interference between structural elements of the        spine, bones, muscle, nerves, is avoided.    -   4. The load bearing capacity of the annulus is directly        dependent upon its ability to develop hoop stress. Accordingly,        inner layers of annulus that are disorganized or fractured        should be treated as part of the nucleus and removed to prevent        tissue extrusion through the annulotomy. The radial dimension of        the space created should be considered r_(n,surg) for use with        the prosthetic selection procedures herein described.    -   5. The disc height, D_(disc), to be used here is the greatest        height reasonably attainable by mechanical distraction,        pressurized injection of an in situ polymerizing disc        prosthetic, patient orientation, or any combination of these.    -   6. The following table is to be used to determine the minimum        modulus necessary to preserve the annulus (Endplate Limit) with        a nucleus prosthesis using surgically derived r_(n), where        r_(d)=1.4 r_(n):

TABLE 1 Removed Implant Nucleus Modulus (MPa) Radius (mm) Free Bonded 716.43 1.83 8 12.58 1.40 9 9.94 1.10 10 8.05 0.89 11 6.65 0.74 12 5.590.62 13 4.76 0.53 14 4.11 0.46 15 3.58 0.40 16 3.14 0.35 17 2.79 0.31 182.48 0.28 19 2.23 0.25 20 2.01 0.22

The following table is to be used to determine the minimum modulusnecessary to preserve the annulus (Endplate Limit) of nucleus prosthesisusing r_(n), where r_(d)=14:

TABLE 2 Removed Implant Nucleus Modulus (MPa) Radius (mm) Bonded Free 71.20 10.8 8 1.09 9.8 9 0.99 8.9 10 0.89 8.0 11 0.81 7.3 12 0.73 6.6 130.66 6.0 14 0.60 5.4

-   -   7. The nucleus Extrusion Limit, P_(n)<P_(ave,defect), must be        applied by measuring the annulus defect diameter and using the        table below to determine whether prosthetic bonding is        necessary. See 8 if the disc height is more than 5 mm.

TABLE 3 Maximum Static Loads for Free Prosthetic Implant Modulus DefectDiameter (mm) (MPa) 0.5 mm 1.0 mm 2.5 mm 3.5 mm 4.9 mm 0.25 2467.451262.45 355.7 111.7 19.1 0.5 4934.9 2524.9 711.4 223.4 38.2 1 9869.85049.8 1422.8 446.8 76.4 1.5 14804.7 7574.7 2134.2 670.2 114.6 2 19739.610099.6 2845.6 893.6 152.8 2.5 24674.5 12624.5 3557 1117 191 3 29609.415149.4 4268.4 1340.4 229.2 3.5 34544.3 17674.3 4979.8 1563.8 267.4 439479.2 20199.2 5691.2 1787.2 305.6 4.5 44414.1 22724.1 6402.6 2010.6343.8 5 49349 25249 7114 2234 382 5.5 54283.9 27773.9 7825.4 2457.4420.2 6 59218.8 30298.8 8536.8 2680.8 458.4

TABLE 4 Maximum Static Loads for Bonded Prosthetic Implant ModulusDefect Diameter (mm) (MPa) 0.5 mm 1.0 mm 2.5 mm 3.5 mm 4.9 mm 0.2522207.05 11362.05 3201.3 1005.3 171.9 0.5 44414.1 22724.1 6402.6 2010.6343.8 1 88828.2 45448.2 12805.2 4021.2 687.6 1.5 133242.3 68172.319207.8 6031.8 1031.4 2 177656.4 90896.4 25610.4 8042.4 1375.2 2.5222070.5 113620.5 32013 10053 1719 3 266484.6 136344.6 38415.6 12063.62062.8 3.5 310898.7 159068.7 44818.2 14074.2 2406.6 4 355312.8 181792.851220.8 16084.8 2750.4 4.5 399726.9 204516.9 57623.4 18095.4 3094.2 5444141 227241 64026 20106 3438 5.5 488555.1 249965.1 70428.6 22116.63781.8 6 532969.2 272689.2 76831.2 24127.2 4125.6

-   -   8. If the surgeon elects to increase the disc height from ITS        diseased height by some form of distraction, then the modulus        required to satisfy the Endplate Limitation decreases by the        factor D_(corrected)/D_(disease). The modulus required to        satisfy the Extrusion Limit changes as shown in FIG. 34, where        P_(ave)(disc height=5 mm)=1. FIG. 34 is represented by Table 5.

TABLE 5 Defect Disc Height (mm) Diameter (mm) 6 mm 7 mm 8 mm 9 mm 10 mm11 mm 12 mm 0.5 0.875035 0.781545 0.708979 0.650957 0.603432 0.5637260.53 1 0.924871 0.860423 0.805617 0.758783 0.718403 0.683248 0.65235 1.50.97868 0.947851 0.914914 0.882715 0.85231 0.824012 0.797824 2 1.0414221.052389 1.048177 1.036207 1.020305 1.002529 0.984021 2.5 1.1196191.185582 1.220976 1.238102 1.243916 1.24254 1.236519 3 1.225811 1.3696621.463281 1.524658 1.564563 1.589739 1.604568 3.5 1.389185 1.6564031.844834 1.980157 2.078436 2.150148 2.202366 4 1.697321 2.2013122.575099 2.857635 3.074185 3.241797 3.372348 4.5 2.588091 3.7822844.702029 5.422998 5.995734 6.455359 6.827046 4.9 9.604787 16.2486821.49539 25.70658 29.12825 31.93488 34.25413

Exemplary Prepolymers

Suitable prepolymers form non-absorbable hydrogels. Examples of suitablenonabsorbable hydrogel compositions are described in U.S. Pat. No.6,296,607. Other compositions that have the appropriate strength, andthat bond to tissue when required to obtain appropriate mechanicalproperties are also suitable.

Non-Absorbable Prepolymers

Prepolymers of polyurethanes can be used as hydrogels. They are formedby encapping triols, or triolized diols with diisocyanate and thenreacting these with excess quantities of water. When the polyolcomponent contains approximately 75% polyethylene oxide and 25%polypropylene oxide the resulting hydrogel can contain up to 90% waterand achieve desirable stability and strength characteristics.

Exemplary prepolymers are the product of reacting about 20% by weight toabout 40% by weight TDI, 65% by weight to about 85% by weight diol andabout 0.5% by weight to about 2% by weight TMP. In one embodiment, thecomposition is the product of reacting in weight ratios about 20% toabout 25% TDI, 70% to about 80% diol and about 0.7% to about 1.2% TMP.In another embodiment, the composition is the result of reacting about23% to about 25% TDI, about 73% to about 77% diol and about 0.7% toabout 1.0% TMP. In yet another embodiment, the composition is the resultof reacting about 24% TDI, 75% diol and about 0.7% to 1.0% TMP. In yetanother embodiment, the diol is 75% polyethylene glycol and 25%polypropylene glycol.

Other suitable compositions are the product of reacting about 20% byweight to about 40% by weight IPDI, 65% by weight to about 85% by weightdiol and about 1% by weight to about 10% by weight TMP. In oneembodiment, the composition is the product of reacting in weight ratiosabout 25% to about 35% IPDI, 70% to about 80% diol and about 2% to about8% TMP. In another embodiment, the composition is the result of reactingabout 25% to about 30% IPDI, about 70% to about 75% diol and about 1% toabout 8% TMP. In another embodiment, the composition is the result ofreacting about 25% IPDI, 70% diol and about 1% to 2% TMP. In yet anotherembodiment, the diol is 75% polyethylene glycol and 25% polypropyleneglycol.

Hydrogels can formed by mixing the above prepolymers with up to 90%water by volume, e.g., 50% water by volume.

Animal Studies

Synthesis of the Implant

Seven hundred grams of Diol UCON 75-H-1400 (Dow Chemical) is heat to 49°C. and stirred under a continuous flow of argon for 24 hours. Theprepared diol is cooled to room temperature (22° C.) and 113.40 g ofToluene Diisocyanate added. The mixture is stirred under an argonblanket and the temperature of the solution increased linearly to60+/−2° C. over a 2 hour period. The mixture is maintained at thesecondition until the % NCO drops to 2.95%. When this target is reached6.26 g of Trimethylolpropane is added, and the mixture stirred underargon at 60+/−2° C. until the % NCO=2.21.

The composition above was use directly (100% polymer) and mixed withequal parts by volume of water (50:50 polymer). These compositions wereinjected into an isolate lumbar segment of a 200 lb pig through a 2.5 mmdiameter annulotomy. FIG. 35 illustrates disc compliance for the naturaldisc, the natural disc with nuclectomy, and the same disc filled with100% and 50:50 versions of the polymer. FIG. 36 illustrates the loadbearing capability of one embodiment of an implant via a plot of % totalload (y-axis) versus displacement (x-axis).

Treatment of Thin Discs by Trans-Axial Approach

FIG. 37 shows common pathologies of spinal discs located in the lumbarregion. Lumbar discs have the largest radius, r_(d) (9-14 mm), and discheight, D_(disc) (9-12 mm), of all the spinal discs. In the diseasedstate [the] A disc can be highly compressed, and in this example thedisc is compressed from a normal 9 mm to 3 mm. Otherwise, the disc doesnot show signs of annulus rupture.

In this case, the absence of bulging or nucleus leakage mitigates theneed for an annulotomy. There are two options: 1) access the nuclearspace via a trans vertebral body approach or 2) inject the prostheticposterior laterally with a small gauge (18 G) needle through theannulus.

The first approach will be considered in this example. The trans-axialapproach comprises the steps of 1) forming a passage through thevertebral body endplates, 2) inserting a coring device into the nuclearspace that sweeps out the nucleus in a radius r_(n)=12 mm, 3)mechanically distracting the endplates to achieve an 8 mm disc height,and 4) filling the space formed with a nonabsorable in situ polymerizingagent.

Since the nucleus is completely removed in this procedure, it isbeneficial to use a nucleus prosthetic that can be delivered as a liquidand can bond to the endplates. Any polymer disclosed herein cansufficiently bond the endplates in this application. Referring now toTable 1, the Endplate Limit specification for the implant modulus is 731kPa. The target disc is 8 mm, so the value obtained from Table 1 must bedecreased as specified in Step 9, 731 kPa×5/8=457 kPa.

Treatment of Thin Discs by Trans-Annulus Approach

If the annulus is in tact, the surgeon may elect not to create a defectin it in order to perform a nuclectomy. In this case, the liquidprosthetic is injected over the existing nucleus. The goal is to restoresome disc height, so a pressurized injection will be used to helpdistract and restore disc height. The nuclear material is in an unknownstate and likely highly fractured and disordered. It acts as a loosecoating on the endplates preventing bonding of the prosthetic to a rigidstructure. Thus, the surgeon will use the column in Table 1corresponding to a free prosthetic.

This section of Table 1 requires high prosthetic modulus, and thesurgeon recognizes that the patient is elderly and has reduced bonedensity. In this case, the Endplate Limit should be strictly respected.Since the annulus will not be damaged by the procedure, the surgeonelects to distract the space to 12 mm. This reduces the modulus from6.57 MPa to 6.57×5/12=2.7 MPa. Using a lower modulus prosthetic willshift load to the annulus, this implant modulus represents a balancebetween stresses on the annulus and those on the endplates.

Treatment for Black Disc

FIG. 37 also depicts common pathologies of the mid spinal region. Onedisc depicts nucleus dehydration and or degeneration. Disc degenerationresults in irritating and often painful degradation products, which mustbe removed. The current standard of care is removal of the nuclearmaterial without replacement. This will result in an eventual loss of50% disc height, causing the bony structures of the spine to closearound nerves and cause pain.

The treatment consists of a 2.5 mm annulotomy made in the annulus,followed by removal of all the degenerated material. Once the nuclearspace is open to atmospheric pressure, the disc height is more easilyincreased by positioning the patient. The target disc height is 9 mm.Due to the highly diseased state of the nucleus, some of the annulus isalso to be removed, and imaging reveals about r_(n)=12 of the r_(d)=14mm of the disc is to be removed. Then from Table 2 (bonded) we obtain aprosthetic modulus of 730 kPa, and this is to be reduced by 5/9 to yield406 kPa. This is the Endplate Limit.

Looking now to Table 4, for free prosthetic, the minimum modulus is 1MPa. From Table 5, we see there is marginal benefit in increasing thedisc height from an Extrusion Limit perspective. In this case, after theheight adjustment is made, the Extrusion Limit dictates the choice ofprosthetic modulus.

Treatment for Bulging Disc

An aneurysm of the disc is a bulge that impinges on nerves, causingpain. Because this represents a weak portion of the annulus, merelyfilling the nucleus to a higher height will not draw the aneurysm inbecause once the load is re-established the weak portion will bulge tothe same extent, since the nuclear pressure under load is notsignificantly improved. This is a portion of the annulus where E_(hoop)is small compared to the rest of the annulus, and therefore can beremoved without loss of annulus energy storage capacity. It must beremoved, if pain is to be alleviated.

In this case, the side of the annulus defect is governed by the extentof the aneurysm and not the size of the surgeon's tools. In this casethe defect will be 4.9 mm in diameter. Given the size of the defect,restoring disc height significantly decreases the modulus requirement.Here, the target disc height will be 9 mm.

As is the case with any defect, the nucleus must be removed. Oneadvantage of a large annulotomy is that the nucleus can be thoroughlyremoved, and the prosthetic can be effectively bonded to the endplates.

Imaging reveals the nucleus is relatively healthy with r_(n)=10 andr_(d)=14. In this case, the surgeon will want to avoid removing andhealthy annulus material. From Table 1, bonded, the Endplate Limit is893 kPa reduced by the disc height improvement by 5/9 to 496 kPa. FromTable 4 the Extrusion Limit is between 1.5-2 MPa. From Table 5, theincrease in disc height decreases the Extrusion Limit by 1/7, and yieldsa modulus of 240 kPa. In this case the Endplate Limit determines thechoice of prosthetic.

Treatment for Permanently Compressed Disc

In this case the achievable distraction is only 5 mm. Here the extrusionminimum is likely to dominate. The surgeon may elect to try to minimizethe annulotomy and use the free condition of Table 1. Table 1 gives 8.4MPa. Given a small annulotomy, the Endplate Limit now dominates. At thishardness, the prosthetic is nearly as hard as the endplates. This is aconsequence of the shortness of the disc height, where furthershortening may cause pain and a normal 20% deflection is a relativelysmall distance.

Decreasing the prosthetic modulus below the Endplate Limit riskspain-inducing deflections under load, especially since the procedureinvolves creating a defect in the annulus.

One option is to enlarge the nuclear space using Table 2, whiledecreasing the annulotomy. Once the annulotomy has been made andmeasured, then the surgeon can determine the amount of nucleus toremove. In this case, the surgeon is approximating a total discprosthetic.

Given the disc's compressed state, the surgeon ends up with a 3.5 mmdefect. Looking on Table 3, this gives an extrusion limit of 6 MPa. Nowturning to Table 2, r_(n) must be enlarged from its present 10 mm to 13mm to get the Endplate Limit down to 6 MPa.

Generalized Indicators for Nucleus Prosthesis Modulus

Table 6 provides general guidelines for choosing a disc prosthesis basedon the state of the disc and achievable end points.

TABLE 6 Treatment Paradigms for Nucleus Replacement Disc Height IncreaseIndication Bonded Restored Nucleus Modulus Procedure Annulotomy ThinDisc Yes Yes No 0.5 MPa   Trans-axial No Lumbar Thin Disc Yes No No 1MPa Trans-axial No Lumbar Thin Disc No Yes No 3 MPa Trans- No Lumbarannulus Thin Disc No No No 6 MPa Trans- No Lumbar annulus Black Disc YesEither Yes 1 MPa annulotomy <2.5 mm Bulging Disc Either No No 2 MPaAnnulotomy <4.9 mm Bulging Disc Yes No No 1 MPa Annulotomy <4.9 mmBulging Disc Yes Yes No 0.25 MPa   Annulotomy <4.9 mm Permanently No NoNo 8 MPa Annulotomy Any size compressed Permanently No No Yes 6 MPaAnnulotomy <3.5 mm compressed Permanently Yes No No 1 MPa Annulotomy<3.5 mm compressed

Exemplary embodiments are provided below.

1. A polymerizable spine disk repair implant, wherein said implant isselected to have a modulus, after polymerization, that distributes thestress caused by a load on the spine evenly between the annulus and theend plates.2. The implant of claim 1 where the stress is distributed in a rangebetween about 3:1 to 1:3 between the annulus and the end plates.3. An implant according to claim 1 in which the implant adheres to thewalls and endplates of the defect into which it is implanted.4. An implant according to claim 1 in which the implant does not adhereto the walls or endplates of the defect into which it is implanted.5. The implant of claim 1 wherein the modulus of the implant is selectedfrom any of Tables 1, 2, 3 or 4 of this application, according to thesizes of various features of the disk being prepared.6. A method for selection of a prosthesis material for repair of aspinal disc, the method comprising:

determining the radius of the disc;

determining the radius of the space that will be filled by theprosthesis;

selecting whether the implant will adhere to the disc and end plates, ornot;

and selecting a defect radius for a defect formed while removing nucleusfrom said disc; and then

using a table having the same information as Table 1 or Table 2 of thisapplication to determine the required modulus for the material used forrepair, after it is cured;

and use a table having the same information as Table 3 or Table 4 ofthis application to ensure that the repair material of the selectedmodulus has a value of maximum static load that is greater than acriterion, at the selected defect radius;

and if so, to use the selected modulus material as the repair material;

and if not, to use a table equivalent to table 3 or table 4 to select amaterial of a higher modulus that will be greater than the criterion.

7. The method of claim 6 where the criterion is a maximum static loadthat is greater than 1000 Newtons.8. The method of claim 6 where a selected material is provided bydilution of a stock material with a physiologically-compatible solution.9. The method of claim 6 wherein the modulus is adjusted for disc heightas a function of defect radius by multiplying it by a factor from atable containing the information of table 5 of this application.10. The method of claim 6 wherein all of the parameters of tables 1-5are contained in a computer program which will calculate the requiredmodulus upon entry of the values selected into said computer program.11. The method of claim 10 where a required disk height adjustment asdescribed in claim 14 is also entered into said program and its effectsincluded in the calculation.12. The method of claim 10 wherein the output of the program is aninstruction of how much to dilute a stock material with aphysiologically-compatible solution.13. An implant for repair of a nucleus of a spinal disk wherein thephysical properties of the annulus are selected to conform to theEndplate Limit, wherein the pressure in the annulus is selected to beequal to the hoop stress, as described herein.14. An implant for repair of a nucleus of a spinal disk wherein thematerial properties of the implant are selected so that after repair,about 25% of the load is carried by the annulus and about 75% of theload is carried by the nucleus.15. An implant according to claim 13 or 14 wherein the implantation ofthe implant is accompanied by an adjustment of disk height to a greatervalue.16. An implant according to claim 15 wherein the implant adheres toadjacent tissue.17. A kit for spine repair, the kit comprising a polymerizing materialfor replacing at least a part of a disc nucleus, a device for deliveringthe polymerizing material, and one or more tables for selection of thedegree of dilution of the polymerizing material according to theconditions required to produce a limitation of the load on the annulusto about 25% of the total load.18. A procedure for nucleus repair, wherein the procedure contains astep requiring the use of a material with the ability to adheresufficiently to surrounding tissue to lower exterior pressure by afactor of 3 or more.

1. A spinal disc implant comprising a foam adapted to completely orpartially replace a nucleus pulposus within a disc nucleus space, thefoam being a nonabsorbable, closed cell and having a Poisson ratio ofless than 0.5.
 2. The implant of claim 1, wherein the foam comprises 30to 50% by volume of closed cell bubbles.
 3. The implant of claim 2,wherein the closed cell bubbles contain less than 50% of a liquid within1 week or less after implantation.
 4. The implant of any precedingclaim, wherein the foam further comprises a radio-opaque marker.
 5. Theimplant of claim 4 wherein the marker is selected from tantalum andbarium sulfate.
 6. The implant of any preceding claim, wherein the foamadheres to at least one of the annulus fibrosus, walls, and endplates ofthe spinal disc cavity.
 7. The implant of claim 6, wherein the foamadherence takes the form of a covalent bond.
 8. The implant of claim 7,wherein the bond strength ranges from 4 lbs/in² to 24 lb/in².
 9. Theimplant of any preceding claim, wherein the foam has a modulus rangingfrom 0.5 to 10 MPa.
 10. The implant of any preceding claim, wherein thefoam has a modulus ranging from 0.5 to 5 MPa.
 11. The implant of anypreceding claim, wherein the foam comprises a polyurethane.
 12. Theimplant of any preceding claim, wherein the foam is sufficientlyhydrophilic to wet tissue surfaces.
 13. The implant of any precedingclaim wherein the foam has a bimodular compliance.
 14. The implant ofany preceding claim, wherein the foam is prepared from: a polyurethaneprepolymer comprising a polymeric polyol end-capped with diisocayanate,and a low molecular weight polyisocyanate.
 15. The implant of claim 14,wherein the polymeric polyol comprises polyethylene oxide andpolypropylene oxide.
 16. The implant of claim 15, wherein the polymericpolyol comprises polyethylene oxide in an amount ranging from 70 to 90%by weight and polypropylene oxide in an amount ranging from 10 to 30% byweight.
 17. The implant of claim 15, wherein the polymeric polyolcomprises 75% polyethylene oxide and 25% polypropylene oxide.
 18. Theimplant of any one of claims 14-17, wherein the polyurethane prepolymeris a trifunctional polyol capped with diisocyanate, the trifunctionalpolyol being formed by trimerizing polymeric diols with trimethylolpropane.
 19. The implant of claim any one of claims 14-18, thepolyurethane prepolymer has a molecular weight ranging from 4500 D to5500 D.
 20. any one of claims 14-19, the low molecular weightpolyisocyanate has a molecular weight of 1000 D or less.
 21. A method ofrepairing a defect in a spinal disc space, comprising: inserting anonabsorbable, closed cell foam having a Poisson ratio of less than 0.5into the defect.
 22. The method of claim 21, wherein prior to theinserting, the method further comprises removing some or all of thenucleus pulposus within the spinal disc space, and the inserting resultsin replacement of the removed nucleus pulposus with the foam.
 23. Themethod of claim 22, wherein the removing further comprises removingportions of the annulus fibrosus in the vicinity of the nucleuspulposus.
 24. A method of repairing a defect in a spinal disc space,comprising: inserting a composition in the area of the defect, thecomposition comprising: (a) a prepolymer, and (b) a foaming component;and curing the composition to form a nonabsorbable, closed cell foamhaving a Poisson ratio of less than 0.5.
 25. The method of claim 24,wherein the cured foam expands to fill the spinal disc space.
 26. Themethod of claim 24 or 25, wherein prior to the inserting, the methodfurther comprises removing some or all of the nucleus pulposus withinthe spinal disc space, and the inserting results in replacement of theremoved nucleus pulposus with the foam.
 27. The method of any one ofclaims 24-26, wherein the cured foam comprises 30 to 50% by volume ofclosed cell bubbles.
 28. The method of any one of claims 24-27, whereinthe inserting comprises injecting the composition.
 29. The method of anyone of claims 24-28, wherein the composition has a viscosity rangingfrom 100 cp to 1000 cp.
 30. The method of any one of claims 24-29,wherein the curing occurs over a period of time ranging from 2 to 5minutes.
 31. The method of any one of claims 24-30, wherein the curedfoam does not swell more than 30% by volume.
 32. The method of any oneof claims 24-31, wherein the foaming component comprises an aqueoussolution.
 33. The method of claim 32, wherein prior to the inserting,the composition is adjusted by mixing the prepolymer with the aqueoussolution in a ratio that achieves a desired foam compliance.
 34. Themethod of claim 33, wherein the prepolymer and aqueous solution aremixed in a volumetric ratio of 1:1.
 35. The method of any one of claims24-34, wherein the prepolymer comprises free isocyanate groups and freeamine groups.
 36. The method of claim 35, wherein the composition curesthrough the interaction of isocyanate groups with amine groups, some ofwhich are also present in the tissue of the implantation site.
 37. Themethod of any one of claims 24-36, wherein the cured foam exchangeswater with the surrounding tissue.
 38. The method of any one of claims24-37, wherein the cured foam is a permanent implant.
 39. The method ofany one of claims 24-38, wherein the inserting further comprises a firstand second application of the composition, wherein the secondapplication bonds to the first application of the cured foam.
 40. Themethod of any one of claims 24-39, wherein the cured foam can besurgically removed from the implantation site.
 41. The method of any oneof claims 24-40, wherein the cured foam provides clinically significantreinforcement to the annulus fibrosus.
 42. The method of any one ofclaims 24-41, wherein the cured foam changes height in response tospinal loads.
 43. The method of any one of claims 24-42, wherein thecured foam begins with a Poisson ratio less than 0.5, changes height inresponse to spinal loads, and subsequently attains a Poisson ratio ofapproximately 0.5.
 44. The method of claim 43, wherein the cured foamhas a Poisson ratio of less than 0.5 at least within 5 days afterimplantation.
 45. The method of any one of claims 24-44, wherein thecured foam translates axial forces originating at the vertebralendplates into radial forces applied to the disc annulus.
 46. The methodof any one of claims 24-45, wherein gas in the foam is replaced withfluids from surrounding tissues within about one month or less.
 47. Themethod of any one of claims 24-46, wherein a bond strength between theprosthetic and tissue ranges from 4 lbs/in² to 24 lb/in².